Optimized pulsatile-flow ventricular-assist device and total artificial heart

ABSTRACT

A method of optimizing a mechanical cardiac pumping device includes modeling the circulatory system of the patient who will receive the mechanical cardiac pumping device and identifying an operating condition of the native heart to which the device will respond. The model is used to determine the required blood volume to be ejected from the device and an initial estimate of the power required to be provided to the mechanical cardiac pumping device is provided in order to provide the required ejected blood volume. The resultant ejected blood volume is evaluated with data obtained from the model and the estimate of the power requirement is then updated. The above steps are iteratively performed until the power required to obtain the necessary ejected blood volume is identified. Possible variations of power and pumping rate that allow the mechanical cardiac pumping device to provide the required volume are determined and the variation that best matches the physiological constraints of the patient and minimizes the power required by the mechanical cardiac pumping device is selected. The steps are iteratively performed until the mechanical cardiac pumping device is optimized to respond to each desired operating condition of the native heart.

BACKGROUND OF THE INVENTION

[0001] 1. Field of the Invention

[0002] This invention relates generally to the field of mechanicalcardiac pumping devices, and, more particularly, to a ventricular assistdevice (VAD) and a total artificial heart (TAH) device and method ofusing same. More specifically, this invention relates to a VAD and a TAHthat are optimized by the new method to produce customized pulsatileblood flow mimicking that of the healthy native heart for eachindividual patient case.

[0003] 2. Description of Related Art

Introduction

[0004] Some medical studies indicate: a) 400,000 new cases of congestiveheart failure are diagnosed annually in the United States; b) amortality rate of 75 percent in men and 62 percent in women; c) standardmedical therapies benefit only a limited percentage of patients withventricular dysfunction; and d) from 17,000 to 66,000 patients per year,in the United States alone, may benefit from a permanent implantableblood pump. Presently, potential cardiac transplant recipients withhemodynamic compromise (inadequate perfusion of the systemic circulationby the native heart) sometimes receive temporary mechanical circulatorysupport as a “bridge” to permit them to survive until cardiactransplantation is possible. It is foreseen that some day mechanicalblood pumps will provide a cost-effective alternative to either cardiactransplantation or long term medical management of patients. It is tothis end that the devices and methods described herein have beendeveloped.

[0005] It is to be understood that for purposes of this document a“ventricular-assist device (VAD)” is a mechanical blood pump thatassists a diseased native heart to circulate blood in the body, and a“total artificial heart (TAH)” is another type of mechanical blood pumpthat replaces the native heart and provides all of the blood pumpingaction in the body.

[0006] In order for a VAD to function optimally, it must both complementthe diseased native heart and make the combined output of the VAD andnative diseased heart emulate the pumping action of the natural healthyhuman heart. That is, it should provide pulsatile flow similar to thatof the healthy heart. In order for a TAH to function optimally, it mustmimic the pulsatile pumping action of the natural healthy human heart.In either case, the device must be sized such that it fits within therequired areas in the patient's body. In order to minimize the size ofthe power supply portion of the device, each device (VAD or TAH) mustuse as little energy and as little power as possible to accomplish therequired function. Thus, there is a need for bio-emulating efficientpump (BEEP) systems for VAD and TAH applications.

[0007] It is known that VADs can be implanted to assist a functioningheart that does not have adequate pumping capability. Often, however,residual cardiac function is not taken into account in the design ofsuch devices, resulting in less than optimal effects. What is needed isa bio-emulating efficient pump (BEEP) system, which works in concertwith the native human heart. The new VAD device and system andoptimization procedure described herein utilize patient specificinformation concerning residual cardiac output to optimize the pumpingaction provided for each individual patient, thereby providing such aBEEP system. The TAH device and optimization procedure described in thisdocument optimize the pumping function provided for each individualpatient, thereby providing such a BEEP system which is customized foreach such patient.

Known Heart Pump Devices

[0008] Previously, a number of devices were developed for blood pumping.Highly specialized pumps have been used to completely replace abiological heart which has been surgically removed. Such known heartpumps may be temporary, or permanently implantable. Temporary heart pumpdevices usually involve either: 1) an attempt to augment a compromisednative heart while it recovers from surgery or some other short-termproblem; or 2) use of the device as a “bridge” to extend the life of apatient by temporarily replacing the native heart until a suitable donorheart can be found for cardiac transplantation.

[0009] Many types of permanently implantable heart pumps have beenproposed and several have been developed. Because the left ventricle ofthe heart, which pumps blood to the entire body except for the lungs,becomes diseased far more commonly than the right ventricle (which pumpsblood only to the lungs), most heart pumps have been developed to assistor replace the left ventricle. Fewer pumps have been proposed, tested,and used for bi-ventricular function (i.e. assisting both the left andright ventricles).

[0010] Known mechanical blood pumps can be roughly divided into threemajor categories: a. pulsatile sacks; b. reciprocating piston-typepumps; and c. pumps with axial or centrifugal impellers. Each categoryhas distinct advantages and disadvantages.

a. Pulsatile Sacks

[0011] Pulsatile sack devices are the most widely tested and usedimplantable blood pumps. These devices employ flexible sacks ordiaphragms which are compressed and released in a periodic manner tocause pulsatile flow of blood. Sack or diaphragm pumps are subject tofatigue failure of compliant elements. They are generally used astemporary heart-assist devices, and they are mechanically andfunctionally different from the present invention described hereafter.

[0012] The intra-aortic balloon (IAB) counter-pulsation device, apulsatile sack device, is readily available. It is a catheter-mountedintra-vascular device designed to improve the balance between myocardialoxygen supply and demand. The first successful clinical application ofthe balloon was reported by Kantrowitz et al. in 1968. The IAB ispositioned in the thoracic aorta and set to inflate at the dicroticnotch of the atrial pressure waveform when monitoring aortic pressure.The diastolic rise in aortic pressure augments coronary blood flow andmyocardial oxygen supply. The IAB is deflated during the isovolumetricphase of left ventricular contraction. The reduction in the afterloadcomponent of cardiac work decreases peak left ventricular pressure andmyocardial oxygen consumption. These units are not portable and arelimited to in-hospital critical care use only. Use of the IAB is now astandard form of therapy for a variety of patients with cardiovasculardisease, primarily reserved for patients with deteriorating heartfunction while awaiting revascularization procedure. In 1993, nearly100,000 IABs were inserted in the United States alone.

[0013] Another example of a pulsatile sack device is the Abiomed™ BVS®device (Abiomed, Inc., Boston, Mass.). It is an externally placeddual-chamber device that is capable of providing short termuniventricular or biventricular support. It has pneumatically drivenpolyurethane blood sacks and it is not intended for long-term support.Also, U.S. Pat. No.4,888,011 to Kung and Singh discloses a hydraulicallydriven dual-sack system; and U.S. Pat. No. 5,743,845 to Runge disclosesa sack-operated bi-ventricular assist device that balances the flow inthe left and right side of the circulatory system.

b. Reciprocating Piston-Type Pumps

[0014] Several types of implantable blood pumps containing a piston-likemember have been proposed to provide a mechanical device for augmentingor totally replacing the blood pumping action of a damaged or diseasedheart. For example, the HeartMate® (Thermo Cardiosystems, Inc., Woburn,Mass.) is a pneumatically powered device that is implanted in the leftupper quadrant of the abdomen. A pneumatic air hose exits from the lowerhalf of the abdominal wall and is attached to a pneumatic power unit.Blood from the cannulated left ventricular apex empties into a pump, atwhich point an external control system triggers pumping. The bloodchamber is pressurized by a pusher plate forcing a flexible plasticdiaphragm upward. This motion propels the blood through an outflowconduit grafted into the aorta, the main artery supplying the body withblood. This device is unique in that the textured, blood-containingsurface promotes the formation of a stable neointima, hence fullanticoagulation is not necessary, only anti-platelet agents arerequired. This device is designed for left ventricular support only. Ituses trileaflet polyurethane valves. There is an electrically poweredversion with percutaneous electric leads connecting the pump to externalbatteries.

[0015] The Thoratec® VAD (Thoratec Laboratories, Pleasanton, Calif.) isa pneumatically powered device that is placed externally on the anteriorabdominal wall. Cannulas pass through the chest wall in a manner similarto that of a conventional chest tube. The device takes blood from theleft ventricular apex and returns it to the aorta. Full systemicanticoagulation is required with this device. It can be used to supporteither ventricle and uses tilting disc type mechanical valves.

[0016] Novacor® (Cedex, France) produces an electrically driven devicethat is implanted in the left upper quadrant of the abdomen and theelectric line and vent tube are passed through the lower anteriorabdominal wall. This system also incorporates a polyurethane blood sacthat is compressed by dual symmetrically opposed pusher plates. Blood istaken from the left ventricular apex and returned to the aorta. Fullanticoagulation is required.

[0017] U.S. Pat. No. 3,842,440 to Karlson discloses an implantablelinear motor prosthetic heart and control system containing a pump witha piston-like member which reciprocates in a magnetic field. The pistonincludes a compressible chamber in the prosthetic heart whichcommunicates with the vein or aorta.

[0018] U.S. Pat. Nos. 3,911,897 and 3,911,898 to Leachman, Jr. discloseheart assist devices controlled in the normal mode of operation tocopulsate and counterpulsate with the heart, respectively, and produce ablood flow waveform corresponding to the blood flow waveform of theassisted heart. The heart assist device is a pump connected seriallybetween the discharge of a heart ventricle and the vascular system. Thispump has cylindrical inlet and discharge pumping chambers of the samediameter and a reciprocating piston in one chamber fixedly connectedwith a reciprocating piston of the other chamber.

[0019] U.S. Pat. No. 4,102,610 to Taboada et al. discloses amagnetically operated constant volume reciprocating pump which can beused as a surgically implantable heart pump or assist. The reciprocatingmember is a piston carrying a check valve positioned in a cylinder.

[0020] U.S. Pat. Nos. 4,210,409 and 4,375,941 to Child disclose a pumpused to assist the pumping action of the heart with a piston movable ina cylindrical casing in response to magnetic forces. A tilting-disk typecheck valve carried by the piston provides for flow of fluid into thecylindrical casing and restricts reverse flow.

[0021] U.S. Pat. No. 4,965,864 to Roth discloses a linear motor usingmultiple coils and a reciprocating element containing permanent magnets,driven by microprocessor-controlled power semiconductors. A plurality ofpermanent magnets is mounted on the reciprocating member. U.S. Pat. No.4,541,787 to DeLong describes a pump configuration wherein a pistoncontaining a permanent magnet is driven in a reciprocating fashion alongthe length of a cylinder by energizing a sequence of coils positionedaround the outside of the cylinder.

[0022] U.S. Pat. No. 4,610,658 to Buchwald et al. discloses animplantable fluid displacement peritoneovenous shunt system. The deviceis a magnetically driven pump, which can be a reciprocating diaphragm,or piston type, or rotary pump.

[0023] U.S. Pat. No. 5,089,017 to Young et al. discloses a drive systemfor artificial hearts and left ventricular assist devices comprising oneor more implantable pumps driven by external electromagnets. The pumputilizes working fluid, such as sulfur hexafluoride to apply pneumaticpressure to increase blood pressure and flow rate.

[0024] Larson et al. in a series of patents (1997-1999, U.S. Pat. Nos.5,879,375; 5,843,129; 5,758,666; 5,722,930; 5,722,429; 5,702,430;5,693,091; 5,676,651; 5,676,162) describe a piston-type pump forventricular assist or total replacement, and associated drivingequipment and power supply. The piston is an artificial heart valve,with valves that have at least two leaflets, acting as a check valve andreciprocating in a cylinder. The walls of the cylinder are a fewmillimeters thick because they contain the coils of a linear electricmotor that must provide pumping power to the VAD. Around the artificialheart valve and inside the cylinder is a hollow cylindrical rare-earthpermanent magnet, which is driven by the linear electric motor. In oneembodiment one device is implanted in series to the aorta (left VAD), oranother device is implanted in series to the pulmonary artery (rightVAD), or two devices are used on both aorta and pulmonary artery(BI-VAD). In a second embodiment one device replaces the left ventricle,or another device replaces the right ventricle, or two devices replacethe whole heart.

[0025] Measurements on experimental devices made with hollow pump coresindicate that such devices are too large to fit in the available spacein the chest cavity in the aorta or pulmonary artery, due to the size ofthe coils necessary to drive the device. For a given volume of bloodpumped per stroke, if the length of the cylinder is restricted such thatthe device fits lengthwise in the human body, then the diameter must beincreased until the. desired volume is reached. The outer diameter ofthe device is severely restricted by the surrounding tissue, and thisleaves little room available in the diameter for the linear magnetmotor. In a bi-ventricular application, if the axes of the two cylindersare located in parallel, then even more space is needed due to thediameters required; and if they are not parallel the magnetic fields ofthe two motors introduce additional electromagnetic losses because thelinear magnet motors are not parallel. Even if the volumetricdisplacement of the device is reduced in order to fit in the availablespace at the expense of throughput, much of the outside diameter of thedevice must still be devoted to the linear motor. However, the mostimportant disadvantage is that the linear motor is driving an annularmagnet containing a one-way valve, so that the ferromagnetic materialcan not be in the core (center) of the motor coils, leading to lowerefficiency.

[0026] At the geometric center (axis) of the motor described by Larsonet al. is the artificial valve acting as the piston, and the blooditself. This structure introduces electromagnetic losses in the devicethat make it less desirable than devices that have ferromagneticmaterial in the geometric center (axis) of the motor coils. In addition,voltage propagates at constant velocity from coil to coil in the linearmagnet motor of the Larson et al. device, and motion of the magnetcarrying the artificial heart valve is coupled to this application ofvoltage, so that the application of current in the Larson et al. deviceis not optimized to minimize the power required to effect theblood-pumping action.

c. Pumps with Axial or Centrifugal Impellers

[0027] After pulsatile devices, rotary pumps, having either centrifugalor axial impellers, are the most widely used and tested devices. Incentrifugal pumps, the blood flow enters axially into a centrifugalimpeller, centrifugal acceleration increases the blood flow velocity,the flow exits radially, and the flow is subsequently decelerated toincrease blood static pressure in the diffusion process. Most suchcentrifugal pumps provide continuous (non-pulsatile) flow; or flow witha small fluctuating pressure trace superimposed on a largersteady-pressure component, such as U.S. Pat. No. 5,928,131 to Prem andU.S. Pat. No. 6,179,773 to Prem and Kolenik.

[0028] Axial pumps direct blood flow along a cylindrical axis, which isin a straight (or nearly straight) line with the direction of the inflowand outflow. The impeller looks like an axial fan, or propeller, insidea nozzle. The impeller imparts acceleration to the fluid, and thesubsequent deceleration (diffusion) process increases the bloodpressure. Most such axial pumps provide continuous (non-pulsatile) flow.

[0029] Some types of axial rotary pumps use impeller blades mounted on acenter axle, which is mounted inside a tubular conduit. As the bladeassembly spins, it functions like a fan or an outboard motor propeller.Another type of axial blood pump, called the “haemopump” uses ascrew-type impeller with a classic screw (also called an Archimedesscrew; also called a helifoil, due to its helical shape and thincross-section). In screw-type axial pumps, the screw spins at very highspeed (up to about 10,000 rpm). The entire haemopump unit is usuallyless than one centimeter (approximately 0.4 inches) in diameter. Thepump can be passed through a peripheral artery into the aorta, throughthe aortic valve, and into the left ventricle. An external motor anddrive unit powers it.

[0030] Axial and centrifugal pumps provide mostly steady (continuous)flow with an imperceptible high-frequency low-amplitude pulsatilecomponent. Various mechanisms have been proposed to convert thispractically steady-flow output into pulsatile flow. However, both axialand centrifugal impeller pumps introduce rapid acceleration anddeceleration forces and large shear stresses in the blood. As is wellknown to those with ordinary skill in the art (Balje, 1981), both typesof turbomachines (axial and centrifugal) are a balanced compromisebetween diameter and speed to provide the specified flow rate andpressure increase. Imposing limits in diameter in order to reduce shearstresses means that the optimum machine requires a higher-speed axialcomponent. Imposing speed limits in order to reduce shear stresses meansthat the optimum machine requires a higher-diameter centrifugalcomponent. It is well know to those with ordinary skill in the art(Wilson and Korakianitis, 1998) that small impellers that can fit insidethe spaces available in the human body will result in high blood shear,due to the high operational speed required.

[0031] The Jarvik 2000® (registered trademark of R. Jarvik, New York,N.Y.) System consists of a small axial flow pump (about the size of aC-cell battery) that is placed in the left ventricular apex and pumpsblood into the aorta. It is still currently being developed and will useexternal batteries and control electronics utilizing induction coils tocarry the control signals through the skin. Power is also deliveredtranscutaneously.

Medical Complications

[0032] According to several medical studies, the above devices aresubject to a number of complications. Insertion of a cannula to feed apump can cause damage to the left ventricle. At least 50 percent ofpatients who are supported for prolonged periods develop infections,including those associated with pneumatic lines or electrical leads.Septic emboli may occur, and the mortality rate is up to 50 percent.VADs may also activate the coagulation cascade, resulting in thrombiformation. This occurs in the approximate range of nine to forty-fourpercent of patients. Stasis of blood within the pump may lead tothrombus deposition. Right ventricular failure may occurperi-operatively with placement of a left VAD. The right heart failurerate may be as high as 33 percent, with one-fifth of those patientsdying from the complication. Rapid recognition of this complication andimplantation of a right VAD may reduce the mortality rate resulting fromright heart failure. Hemorrhage occurs in about 27 to 87 percent ofpatients who require mechanical ventricular assistance. Hemorrhage isalso related to inflow and outflow cannulae and to anticoagulationrequired with the devices.

[0033] One of the most important problems in axial and centrifugalrotary pumps involves the interface between the edges of the blades andthe blood flow. The outer edges of the blades move at high speeds andgenerate high levels of shear. Red blood cells are particularlysusceptible to shear stress damage, as their cell membranes do notinclude a reinforcing cytoskeleton to maintain cell shape. Lysis of redblood cells can result in the release of cell contents and triggersubsequent platelet aggregation. Lysis of white blood cells andplatelets also occurs upon application of high shear stress. Evensublytic shear stress leads to cellular alterations and directactivation and aggregation of platelets. Rotary pumps generally are notwell tolerated by patients for prolonged periods. In medical tests,animals placed on these units for a substantial length of time oftensuffer from strokes, renal failure, and other organ dysfunction. Thedevice and method of optimization disclosed herein minimizes the above,and other, known complications resulting from implantation of either aVAD or a TAH.

Desirable Pump Characteristics

[0034] In many patients with end stage heart disease, there is enoughresidual function left in the native heart to sustain life in asedentary fashion, but insufficient reserve for even minimal activity,such as walking a short distance. This residual function of the diseasednative heart is typically not considered in the design of most VADs.Most known VADs are designed to assume complete circulatoryresponsibility and to receive blood from the cannulated ventricular apexof the particular ventricle they are “assisting,” in what is commonlycalled “fill to empty” mode. It generally takes one or more contractionsof the diseased native ventricle to supply enough blood to the VAD. Oncea pre-specified volume of blood is accumulated in the VAD, then theejection phase of the VAD is initiated. Thus, most known VADs operate inthis “fill-to-empty” mode that is in random association with nativeheart contraction, and can be installed in parallel to the nativeventricle or in series. These constructions are not considered to“complement” the native heart, as does the present invention.

[0035] At least some residual cardiac function is present in themajority of patients who would be candidates for mechanical circulatoryassistance. It is preferable for the natural heart to continuecontributing to the cardiac output even after a mechanical circulatorydevice is installed. This points away from the use of total cardiacreplacements and suggests the use of assist devices whenever possible.However, the use of assist devices also poses a very difficult problem.In patients suffering from severe heart disease, temporary orintermittent crises often require artificial pumps to provide bridgingsupport which is sufficient to entirely replace ventricular pumpingcapacity for limited periods of time. Such requirements arise in thehours or days following a heart attack or cardiac arrest, or duringperiods of certain life threatening arrhythmias. Therefore, there is animportant need to provide a pump and method that can meet a widespectrum of requirements by providing two different and distinct pumpingfunctions, assisting the native heart and total substitute pump support.

SUMMARY OF THE INVENTION

[0036] The present invention provides a cardiac ventricular-assistdevice and method of optimizing any design of VAD or TAH wherein theamount of power required by the device is minimized to the extentnecessary to complement the cardiac output of the native heart, and nomore. In this manner, the weight and size of the device are kept withinsuitable reasonable ranges to permit placement of the VAD/TAH within thebody of the subject patient using the new device.

[0037] The present invention further provides a VAD and method whereinthe principles of unsteady thermodynamics and fluid mechanics are usedto provide a uniquely optimized pulsatile blood flow which complementsthe cardiac output of the individual native human heart. It is to beunderstood that throughout this document, when the terms “optimize” and“complement” are used in reference to the devices and systems of thepresent invention, it is meant that at each heart beat and stroke of theVAD (used here to mean either the L-VAD, R-VAD, BI-VAD or TAH asdescribed below), several actions are carefully timed such that:

[0038] a) the native heart is allowed to pump as much blood as it can onits own before the VAD is activated;

[0039] b) as the blood-ejection phase of the native heart nearscompletion, the VAD is energized to provide additional pumping action;

[0040] c) the additional pumping action reduces the back pressure inthat native ventricle so that the native ventricle pumps more than itwould have pumped unaided;

[0041] d) the timing of the action, length of pumping stroke, and rateof pumping (stroke displacement versus time and resulting power inputversus time) of the VAD are related to the native heart ejected bloodvolume and rhythm in a manner that minimizes power input to the VADwhile meeting physiological constraints;

[0042] e) the optimization processes in d) take into account the dynamicinteraction between the native heart and the VAD; and

[0043] f) the optimization process and the control scheme are integratedwith the resulting changes in blood ejected per heart beat and heartrate (beats per minute) by the combined action of the native heart andthe VAD.

[0044] Before turning to the Figures, it is considered useful to providesome introductory material. The present invention, described below, isdistinct from each of the three categories of mechanical circulatorysupport devices previously described, and consolidates the advantagesand avoids the disadvantages of each category. First, it is carefullynoted that several of the devices described in the known art mentionthat the power input is “optimized”, but they do not describe how thisis accomplished. The optimization method described herein can be appliedto all existing VAD and TAH devise that have been devised to date, orwill be devised in the future.

[0045] The pump of the present invention has ferromagnetic material asthe solid center of the motor coils, thus providing a more compactarrangement of the electromagnetic fluxes than pumps withnon-ferromagnetic centers, and simultaneously permitting reduction ofelectromagnetic losses in use. Ultimately this permits placement of adevice that can pump sufficient volume per stroke at the outlet of thenative ventricles and allows the power supply to be smaller than waspossible with previous cardiac pumping devices. The remote hydraulicdrive and power supply/controller assembly are located in the abdomen,thus allowing practically all available space in the vicinity of theheart for use by the device. Power is transmitted hydraulically from theabdomen to the blood pump in the vicinity of the heart. Also,electromagnetic losses are not introduced by the location of the twopumping devices (artificial heart valves) in non-parallel configurationin the vicinity of the aorta and pulmonary artery.

[0046] Details of the dynamics of the pumping action of the human hearthave been incorporated for the first time into the design of the VADsand TAHs in the present invention. Understanding these details:

[0047] is essential for optimization of the timing of unsteady-flowevents in the heart-pumping cycle;

[0048] directly impacts the optimum geometric shape of the artificialdevices; and

[0049] identifies prerequisite means to minimize shear stresses on theblood (reducing blood-cell lysis) and optimizing energy flows (reducingthe power input required to produce the required blood flow and pressurecharacteristics).

[0050] The adult heart is located between the lungs and is about thesize of a large grapefruit, weighing 0.2 to 0.5 kg (0.44 to 1.1 pounds),depending on the size of the individual. The cardiovascular systemperforms two major tasks: it delivers oxygen and nutrients to bodyorgans; and removes waste products of metabolism from tissue cells. Itsmajor components are: the heart (a two-sided biological pump); and thecirculatory system of elastic blood vessels (veins and arteries) thattransport blood. As an example, the heart of a 70 kg (154 pounds) humancirculates about 6 kg (13.2 pounds), or 6 L (6.34 qt.s), of blood.

[0051] The human heart is divided into four chambers: the right atriumand right ventricle; and the left atrium and left ventricle. The wallsof the chambers are made of a special muscle called the myocardium thatcontracts rhythmically under electric stimulation. The left and rightatria are separated from each other by the atrial septum; and the leftand right ventricles are separated from each other by the ventricularseptum.

[0052] In the circulatory system, blood returns by the venous systemfrom the body and enters the heart through the right atrium, thensubsequently blood enters the right ventricle. Each time the rightventricle contracts, it propels this blood (low in oxygen content) intothe lungs, where it is enriched with oxygen. Pulmonary veins returnthe-blood to the left atrium, then subsequently the blood enters theleft ventricle. The left ventricle, which traditionally has beenconsidered as the main pumping instrument of the heart, ejects the bloodthrough the main artery, the aorta, to supply oxygenated blood to thevarious organs of the body. The organs use the oxygen and with capillaryaction between the arterioles and the venules return the blood to thevenous system and the right atrium. The pumping action of the left andright side of the heart generates pulsatile flow and pressure on theaorta and pulmonary artery, discussed further below.

[0053] Blood is kept flowing in this pulsatile cycle by a system of fourone-way valves in the heart, each closing an inlet or outlet in one ofthe heart's four chambers at the appropriate time in the cardiac cycle.The valve system helps maintain a pressure difference between the rightand left sides of the heart. The aortic valve and the pulmonary valveeach have three tissue cusps (leaflet flaps), referred to as “semilunarvalves” because of the crescent shape of these cusps. The tricuspid andmitral valves separate the atria from the ventricles. The mitral valvehas two cusps and the tricuspid valve has three cusps. In addition, thecusps have thin chords of fibrous tissue (chordae tendineae), whichtether the valves to the ventricular walls. When the ventriclescontract, small muscles in their walls (papillary muscles) restrictclosure of the mitral and tricuspid valve leaflets, preventing them fromoverextending.

[0054] Electric currents control the,pumping motion of the heart. Thecurrents originate in the sinus node (the heart's natural pacemaker), amicroscopic bundle of specialized cells located in the superior portionof the atria. The currents travel through a network of conducting fibersto the atrioventricular or AV node, the bundle of His, and the Purkinjefibers. The electric currents cause impulses that are transmitted andpropagate in a wave fashion through the muscle fibers of the left andright atria to the atrioventricular node, located on the juncturebetween the right and left sides of the heart where the right atrium andright ventricle meet. From the atrioventricular node, they travel alongthe bundle of His and the Purkinje fibers through the muscles of theright and left ventricles. Most currents in the heart are less than amillionth of an Ampere, but they exert a powerful influence on the heartmuscle.

[0055] The new VAD utilizes electromagnetic coils to drive ahigh-ferromagnetic-constant driving magnet in a reciprocating fashion soas to act as a piston for hydraulic fluid. The resultant movement ofhydraulic fluid through the system in turn moves another magnet, whichis annular, and which also drives in a reciprocating fashion. Themovement of the driven annular magnet in turn moves still anothermagnet, an annular valve seat magnet, which supports a one-way valve.This valve seat magnet is located inside the annular driven magnet, thetwo magnets sharing a common center axis, hence coupling them together.The one-way valve pushes blood through the ascending aorta of the heartwhen the valve is pushed forward, and allows blood to flow freely pastwhen the one-way valve is moved backward.

[0056] The present invention provides a ventricular-assist device andmethod for optimizing same that can be utilized to assist either theleft ventricle (L-VAD) or right ventricle (R-VAD) of the native humanheart or, if necessary, to assist both cardiac ventricles (BI-VAD). TheL-VAD, R-VAD and BI-VAD devices all utilize principles of unsteady fluidmechanics to provide a uniquely individualized optimized pulsatile bloodflow for each particular patient.

[0057] In an alternative embodiment, a total artificial heart (TAH)device that utilizes the principles of unsteady fluid mechanics providesa uniquely individualized optimized pulsatile blood flow for eachparticular patient. The optimized pulsatile blood flow mimics that ofthe native heart while simultaneously minimizing the power required todrive the TAH device.

[0058] Accordingly, it is among the goals of the present invention toprovide a cardiac pump (VAD or TAH) device and system, and method forcontrolling and operating same which permit customized, optimized“assist” or “total” (“complete”) cardiac pumping support for anindefinite period of time. Under appropriate conditions, the new VADacts synergistically with the native heart to provide a seamlessaugmentation to the otherwise suboptimal output of the diseased nativeheart. This allows the new pump device (VAD) to take advantage of thenatural, non-hemolytic pumping action of the native heart to the fullestextent possible to minimize red blood cell lysis, and to reducemechanical stress on the VAD system pump, requiring less volume, lessenergy, and hence allowing longer pump life and longer battery life.

[0059] Accordingly, in furtherance of the above objects and goods, thepresent invention is, briefly, a method of optimizing a mechanicalcardiac pumping device includes modeling the physical system, orportions thereof, of the patient who will receive the mechanical cardiacpumping device and identifying an operating condition of the nativeheart to which the device will respond. The model is used to determinethe required blood volume to be ejected from the device and an initialestimate of the power required to be provided to the mechanical cardiacpumping device is provided in order to provide the required ejectedblood volume. The resultant ejected blood volume is evaluated with dataobtained from the model and the estimate of the power requirement isthen updated. The above steps are iteratively performed until the powerrequired to obtain the necessary ejected blood volume is identified.Possible variations of power and pumping rate that allow the mechanicalcardiac pumping device to provide the required volume are determined andthe variation that best matches the physiological constraints of thepatient and minimizes the power required by the mechanical cardiacpumping device is selected. The steps are iteratively performed untilthe mechanical cardiac pumping device is optimized to respond to eachdesired operating condition of the native heart.

[0060] The mechanical system for accomplishing the new method is,briefly, a system for assisting cardiac ventricular function, the systemincluding a hydraulic pumping assembly and a cardiac ventricular assistdevice (VAD) in fluid communication with the hydraulic pumping assembly,wherein the hydraulic pumping assembly includes an encapsulatedhydraulic pump having a pumping chamber for retaining hydraulic fluidtherein. The pumping chamber has opposed first and second ends and atleast one electromagnetic coil surrounding the pumping chamber. Asubstantially solid high ferromagnetic-constant magnet is disposedlongitudinally, slideably and reciprocally within the pumping chamber toact as a piston for driving hydraulic fluid within the pumping chamberin response to signals from a battery/controller assembly. A fluid linehas a first end and a second end. The first end of the fluid line isconnected to and in fluid communication with the first end of thepumping chamber and the second end of the fluid line is connected to andin fluid communication with the second end of the pumping chamber. TheVAD is in fluid communication with the fluid line at a point on thefluid line after the point of connection of the check valve and beforethe connection of the second end of the fluid line and the second endpump chamber. A battery/controller assembly is operatively connected tothe check valve and to the at least one electromagnetic coil to provideelectric power and control signals to the pump. The battery controllerassembly is in electrical communication with the native heart of thepatient using the system, to thereby receive signals corresponding tophysiological parameters from the native heart for transfer to the VAD.

[0061] The new VAD device is, briefly, a device to assist the functionof a cardiac ventricle, the device having a first magnet with an opencenter and formed of high ferromagnetic-constant material. A firstvessel of the device surrounds the first magnet and defines a space influid communication with the blood flow output great vessel associatedwith the diseased ventricle of a patient using the device, the firstmagnet being movable within the first vessel in substantiallyfluid-tight relation thereto. A second magnet is formed of highferromagnetic-constant material in magnetic communication with the firstmagnet, so that the magnetic fluxes of the first magnet and the secondmagnet affect each other, so that the first magnet and the second magnetare biased toward and tend to lock to one another, to thereby move inthe same direction as one another. A second vessel encases the secondmagnet and defines a space and is movable within the space insubstantially fluid-tight relation to the second vessel, the space beingdefined by the second vessel being in fluid communication with ahydraulic pump for actuation the second magnet. A one-way valve isconnected to the first magnet, the one-way vale being movable with thefirst magnet, and adapted to cause movement of blood from the diseasedventricle to and into the great vessel associated with the diseasedventricle.

[0062] These and other advantageous features of the present inventionwill be in part apparent and in part pointed out herein below.

BRIEF DESCRIPTION OF THE DRAWINGS

[0063]FIG. 1 is a schematic view generally identifying a bio-emulatingefficient pump (BEEP) system. The Figure specifically illustrates theleft ventricular-assist device (L-VAD) embodiment of a BEEP system atthe beginning of the blood-pumping stroke.

[0064]FIG. 2 is a schematic view of the L-VAD embodiment of a BEEPsystem of FIG. 1, wherein the system is near the middle of theblood-pumping stroke.

[0065]FIG. 3 is a schematic view of the L-VAD embodiment of a BEEPsystem of FIG. 1, wherein the system is at the beginning of the returnstroke.

[0066]FIG. 4 is a schematic view of the L-VAD embodiment of a BEEPsystem of FIG. 1, wherein the system is near the middle of the returnstroke.

[0067]FIG. 5 is a cross-sectional view of the hydraulic pump of the BEEPsystem of FIG. 1, along line 5-5.

[0068]FIG. 6 is a cross-sectional view of the L-VAD of the BEEP systemof FIG. 1, along line 6-6.

[0069]FIG. 7 is a schematic concept illustration of the human heartillustrating the location of an L-VAD in place of at least part of theascending aorta.

[0070]FIG. 8 a schematic sectional view of a human torso O, illustratingthe location of the main components of an L-VAD embodiment of BEEPsystem 35 in the human body. The L-VAD is shown in place of theascending aorta, and the hydraulic pump and battery/controller assemblyare illustrated in the abdominal cavity.

[0071]FIG. 9 is a concept illustration of the human heart illustratinglocation of proximity sensors embedded in the endocardial surface of theleft and right ventricles, and mounted on the pericardium.

[0072]FIG. 10 is a concept illustration of the human heart illustratingthe KG diaphragm in late diastole.

[0073]FIG. 11 is a concept illustration of the human heart illustratingthe KG diaphragm in early systole.

[0074]FIG. 12 is a concept illustration of the human heart illustratingthe KG diaphragm in late systole.

[0075]FIG. 13 is a concept illustration of the human heart illustratingthe KG diaphragm in early diastole.

[0076]FIG. 14 is a graph illustrating typical pressure-volume diagramsof a native healthy heart and a native diseased heart.

[0077]FIG. 15 is a left-ventricle pressure versus time diagram of anative healthy heart and a native diseased heart.

[0078]FIG. 16 is a graph illustrating the relationship between thetravel of the piston of the present device and the residual cardiacoutput provided by the native diseased heart.

[0079]FIG. 17 is a series of graphs comparing the position and powerrequirements of a prior art pumping system and the present BEEP systemwith respect to a typical electro cardio gram (ECG) trace.

[0080]FIG. 18 illustrates the location of three coils of one embodimentof the BEEP system and the corresponding current flow sequence in thecoils.

[0081]FIG. 19 illustrates the location of three electromagnetic coils inone embodiment of the BEEP system and the corresponding current flowsequence in the coils when only two of the coils are used to move thepiston.

[0082]FIG. 20 is a concept illustration of the human heart illustratingthe location of a right ventricular-assist device (R-VAD) embodiment ofthe BEEP system.

[0083]FIG. 21 a schematic view of the human torso illustrating thelocation of the main components of an R-VAD embodiment of a BEEP systemin the human body.

[0084]FIG. 22 is a concept illustration of the human heart illustratingthe location of a bi-ventricular-assist device (BI-VAD) embodiment ofthe BEEP system.

[0085]FIG. 23 a schematic view of the human torso illustrating thelocation of the main components of a BI-VAD embodiment of a BEEP systemin the human body

[0086]FIG. 24 is a concept illustration of a total artificial heart(TAH) embodiment of the BEEP system.

[0087]FIG. 25 a schematic view of the human torso illustrating thelocation of the main components of a TAH embodiment of a BEEP system inthe human body.

[0088]FIG. 26 is a schematic view generally identifying analternative-component configuration of an L-VAD embodiment of a BEEPsystem.

[0089]FIG. 27 is a concept illustration of the human heart illustratingthe design and location of an alternative ejection volume measuringapparatus.

[0090]FIG. 28 is a diagrammatic illustration of the main components ofthe circulation system in the BI-VAD embodiment.

[0091]FIG. 29 is a flow chart schematically illustrating the developmentof the mathematical model (equations 7 and 8) for the dynamic systemincluding the new VAD, in this case the L-VAD.

[0092]FIG. 30 is a flow chart schematically illustrating application ofthe power optimization process in a system including the new ventricularassist device (VAD), in this case the L-VAD.

[0093]FIG. 31 is a flow chart schematically illustrating themulti-input, multi-output control system for performing the new process,and the controller optimization process.

[0094]FIG. 32 is a flow chart schematically illustrating application ofthe new process in a system including the new VAD in an L-VADarrangement.

[0095]FIG. 33 is a flow chart schematically illustrating application ofthe new process in a system including the new VAD in a BI-VADarrangement.

[0096]FIG. 34 is a flow chart schematically illustrating application ofthe new process in an alternative system including the new totalartificial heart (TAH).

DETAILED DESCRIPTION OF THE INVENTION

[0097]FIGS. 1 through 4 are schematic illustrations of the BEEP systemof the present invention, and the structural elements thereof. For theconvenience of the reader, the unique power-optimizing andcontroller-optimizing methods, which are major aspects of the invention,and they are incorporated in the new BEEP system, are illustratedschematically by flow charts in FIGS. 28-34, to be described further,later herein.

[0098] The elements of the new BEEP system, as shown in FIG. 1, forexample, and generally designated 35, compose three primary components:a ventricular assist device (VAD), which in this embodiment is an L-VAD,generally designated 74 (shown on the left side of FIG. 1). L-VAD 74 isactuated by a hydraulic pump, generally designated 42, and controlled bya battery/controller assembly, generally designated 65. It is to beunderstood that the new BEEP system 35 will be referred to throughoutthis document by the same reference numeral, in relation to a variety ofembodiments. Thus, BEEP system 35 may include a L-VAD, R-VAD, BI-VAD, orTAH, all of which are described further herein, or the system mayinclude alternative embodiments of any of the VADs or the TAH describedbelow. The BEEP system is only a vehicle for the other aspects of theinvention, the optimization process described in FIGS. 28-34. Theoptimization process can be applied to any current or future apparatusdesign of L-VAD, R-VAD, BI-VAD or TAH. The BEEP system per se, however,is nonetheless considered to be another important aspect of theinvention, regardless of which embodiments of the various components areincluded in the system. Further in regard to the various embodiments ofthe system, if certain aspects of the overall system are not describedin detail as being different or distinguishable from the otherembodiments, they are considered to be the same or equivalent to thosepreviously or later described.

[0099] BEEP system 35 utilizes electromagnetic coils 46, 48, and 50 todrive a high ferromagnetic-constant solid cylindrical driving magnet 40in reciprocating fashion along the length of hydraulic pump 42. Whilethree such coils are preferred, it is to be understood that the newsystem 35 and the alternative embodiments thereof can operate adequatelywith more than or fewer than three electromagnetic coils on pump 42.Driving magnet 40 is acting as a piston in a hydraulic pump. Theinterior vessel of the hydraulic pump may or may not incorporate endcaps 56 and 57 as part of its hydraulic-vessel design. However, thepresence of the end caps made of ferromagnetic material assist indirecting the flux lines from the surrounding coils to the drivingmagnet 40. It will be obvious to those skilled in the art that severalalternative embodiments can be contemplated by changing the crosssectional areas of the components, which may be circular, rectangular,or a number of other closed shapes.

[0100]FIG. 5 shows pump 42 in cross section and illustrates the externalcylindrical surface of driving magnet 40 mating with the interiorcylindrical surface of electromagnetic coils 46, 48, and 50. Thesesurfaces, whether shaped as the preferred cylinders, or otherwise, arenevertheless sized and shaped to slidingly interact as well as tominimize leakage of hydraulic fluid therebetween. It is understood thatthe function of pump 42 is to use one or more electromagnetic coils todrive one or more magnets in a way to provide motion to driven magnet44; and several alternative embodiments can be used to accomplish thisfunction. It will be obvious to those with ordinary skill in the artthat there can be many variations on the cross-sectional view of thecomponents, on the exact orientation of the electromagnetic fields, onthe exact orientation of the magnets, on the type of hydraulic orpneumatic fluid, and on the details of the design of the vesselcontaining the hydraulic or pneumatic fluid etc, and these alternativeembodiments are herein included. It is understood that severalalternative embodiments to minimize leakage from the high to the lowpressure of the hydraulic fluid and blood, and alternative embodimentsto minimize friction between sliding components, are conceived andconsidered acceptable alternative designs in the present invention.

[0101] End caps 56, 57 are also made of ferromagnetic material, and aredisposed on opposite ends of pump 42. End caps 56, 57 are provided withcentral openings 56 a, 57 a, so that the interior space defined by theelectromagnetic coils is in fluid communication with main hydraulic line60 at each end of the pump cylinder, permitting the hydraulic fluid toflow in and out of the pump cylinder, as described further hereafter.

[0102] End caps 56, 57 also serve to concentrate the magnetic flux ofelectromagnetic coils 46, 48 and 50 in a smaller combined area, therebyimproving pump efficiency. The shape of end caps 56 and 57 assists inthe optimal placement and concentration of magnetic flux lines andminimization of the weight and dimensions of system components. Whileend caps 56, 57 act as “stops” for the piston, there may also beprovided with separate “stops” of known construction, for example, asillustrated in FIG. 19.

[0103] Magnet 40 is preferably entirely solid and thus is sometimesreferred to herein as “solid magnet 40” for convenience of the reader.However, magnet 40 may be only substantially solid; i.e., there could bea small through-hole plugged with plastic, for example, or otherconceivable interruptions to the integrity of the magnet 40 which wouldnot prohibit system 35 from working sufficiently in the present system.However, for most efficient, ideal operation, magnet 40 is entirelysolid.

[0104] Solid magnet 40 acts as a piston to apply force to hydraulicfluid 52 to thereby ultimately move a driven annular magnet 44 (asindicated by arrows A, in FIG. 1) along the length of L-VAD 74. Themagnetic flux of annular driven magnet 44 and annular valve-seat magnet54 are along the axial length of L-VAD 74 so that they are biased towardand tend to lock to one another. Movement of annular driven magnet 44 inturn moves high-ferromagnetic-constant annular valve-seat magnet 54.Blood 78 is therefore pumped by L-VAD 74 in the direction of the flowarrows B through aortic arch 80, as shown in FIG. 1.

[0105] Except for hydraulic fluid leakages, the reciprocating motion ofsolid driving magnet 40 is in phase with the reciprocating motion ofannular driven magnet 44, but is slightly out of phase with thereciprocating motion of valve-seat magnet 54 due to flood, hydraulicfluid, and electromagnetic inertia effects. The out-of-phase separationof driven magnet 44 from valve-seat magnet 54 varies throughout thereciprocating cycle. These delays are accounted for in the time {t}expressions in equation (7) of the new process described later herein.

[0106] The reciprocating movement of driving magnet 40 along the lengthof hydraulic pump 42 is controlled by power, voltage and current frombattery 62 to electromagnetic coils 46, 48 and 50 in the sequencedepicted in FIGS. 17, 18 and 19 and described below. The timing andmagnitude of the current from battery 62 is controlled by controller 64in response to ECG signals initiated from the ECG signal 66, signalsfrom measurements of ejected blood volume, and as a result of theoptimization process explained below. Battery 62 and controller 64 canbe connected as a battery/controller assembly 65, as illustrated, orutilized as separate components. The movement of driving magnet 40 isslightly out of phase with the magnetic field along electromagneticcoils 46, 48 and 50 due to electromagnetic hysteresis effects, which arealso accounted for in the time {t} expression of equation (7) describedherein below. The out-of-phase separation of driving magnet 40 from themagnetic field of electromagnetic coils 46, 48 and 50 varies throughoutthe reciprocating cycle and is mathematically accounted for by theoptimization method described later herein to minimize the powerrequired for operation of the new system.

[0107] Inside an inner sleeve 68 is contained an annular valve-seatmagnet 54 which contains one-way valve 70. The sliding facing surfacesof annular valve-seat magnet 54 and the inner sleeve 68 are sized andshaped to be substantially fluid-tight to minimize leakage of bloodtherebetween. Similarly, the mating surfaces of annular driven magnet 44and the inner sleeve 68 and outer sleeve 72 (shown in FIG. 6) aredesigned to minimize leakage between sliding facing surfaces thereof.

[0108] It is to be understood that several alternative embodiments tominimize leakage from the various mating elements are conceived. It isfurther to be understood that all elements of the new pumping device andthe entire system for operation thereof are formed of suitablebiocompatible, surgical grade materials. Such materials may beappropriately selected from materials that are now known, as well as newmaterials, which may yet be developed.

[0109] Hydraulic pump 42 drives hydraulic fluid 52 in the direction ofthe flow arrows through hydraulic line 59 and into annular space 86,located between inner sleeve 68 and outer sleeve 72. The reciprocatingmotion of hydraulic fluid 52 moves annular driven magnet 44, alsolocated between inner sleeve 68 and outer sleeve 72. By reversing thedirection of current flow in electromagnetic coils 46, 48 and 50, thedirection of driving magnet 40 is reversed, hence the direction ofhydraulic fluid 52 is also reversed; it follows that the direction ofdriven magnet 44 is reversed as well. As annular driven magnet 44 ismoved by the flow of hydraulic fluid 52, magnetic interaction withvalve-seat magnet 54 causes valve-seat magnet 54 to move along withannular driven magnet 44. Because one-way heart valve, for example, asindicated schematically at 70, is secured to valve-seat magnet 54,one-way valve 70 moves in the same direction as valve-seat magnet 54 andannular driven magnet 44. When one-way valve 70 moves in a directionaway from aortic valve 76, it is closed and pushes blood 78 throughaortic arch 80.

[0110] One-way valve 70 can be any artificial or natural heart valve.Some known valves are mechanical, some are biological and some are madewith compliant man-made materials. Some one-way valves may alsoeventually be made with stem cell research. Depending upon theparticular type of valve selected for the one-way valve 70, limits maybe imposed on the optimization process of equation (7), due to thepressure differences the particular valves can withstand (e.g. prolapsemay occur with some compliant heart valves). Such differences are takeninto account in the selection for a particular system as may benecessary.

[0111]FIG. 1 depicts the state of BEEP system 35 at late systole of thenative human heart, when valve-seat magnet 54 is at the beginning of itspumping stroke along the length of L-VAD 74. At the stage of the cycleshown in FIG. 1 the ECG signal 66 and other volume and pressure signalshave been transmitted along wire 63 to controller 64. (Wire 63 may alsobe inside of conductor 410 in the embodiment shown in FIG. 27 anddiscussed hereafter.) In response to these signals and the newoptimization process, controller 64 discharges electrical power, voltageand current to hydraulic pump 42 along wires 67, 69 (and 71, in somecases). Specifically, current from battery 62 has energizedelectromagnetic coils 46 and 48. In response to the energization ofcoils 46 and 48, driving magnet 40 has begun to move away from itsposition within electromagnetic coil 46 and has moved partially withinthe walls of electromagnetic coil 48 (a cross-sectional view of drivingmagnet 40 and electromagnetic coil 46 is shown in FIG. 5).

[0112] Further with reference to FIG. 1, the movement of driving magnet40 has forced hydraulic fluid 52 to move through main hydraulic line 60and secondary hydraulic line 82. The motion of hydraulic fluid 52 placespressure on both driven magnet 44 and check valve 84. Check valve 84 isclosed, as is normally the case, securing the required pressure gradientbetween the high-pressure and low-pressure imposed by the motion ofdriving magnet 40 within hydraulic pump 42. Due to pressure fromhydraulic fluid 52, annular driven magnet 44 has just begun to movealong the length of L-VAD 74, within annular space 86. Magneticinteractions have caused valve-seat magnet 54 to move in a correspondingmanner. Driven magnet 44 is located slightly ahead of valve-seat magnet54 due to electromagnetic and fluid inertia. A cross-sectional view ofL-VAD 74, through outer sleeve 72, annular driven magnet 44, innersleeve 68, and valve-seat magnet 54 is shown in FIG. 6. The function ofthe two magnets, 44 and 54, is to magnetically “lock” to each other sothat the movement of magnet 54 is affected by the movement of magnet 44.By “lock” it is meant that the motion of one magnet affects the motionof the other magnet via their magnetic interaction, even though thedynamics of the system may dictate that the motions of the two magnetsmay be out of phase. Hydraulic vessel (or “sleeve” in some cases) 72 formagnet 54 and blood vessel 68 for magnet 44 may be concentric or not,parallel or not, and may have any cross-section. It will be obvious tothose with ordinary skill in the art that there are several alternativeembodiments for the cross-sectional view of the hydraulic and bloodvessels (parallel axes or not, concentric axes or not, circular,rectangular or other cross section etc) and the exact location andorientation of north and south poles of the magnets, and these areincluded herein.

[0113] Aortic valve 76, located at the outlet of the left ventricle 90,has been retained open by the beginning of the movement of one-way valve70, which is closed and is being moved upward by driven magnet 44. Thedifference in axial location of driven magnet 44 and valve-seat magnet54 is due to fluid inertia, but also due to magnetic inertia. Neitherfluid inertia nor magnetic inertia is accounted for in the prior art.Although in this embodiment it is preferred that one-way valve 70 is anartificial valve of known or newly developed variety, valve 70 may also,if desired or necessary, be a natural heart valve or a one-way valveformed of tissue (human or other animal).

[0114] The movement of closed one-way valve 70 is beginning to pumpblood along the length of the ascending aorta 88 and into the aorticarch 80.

[0115]FIG. 2 depicts the state of BEEP system 35 halfway through thepumping motion of L-VAD 74. In this figure driving magnet 40 has movedwithin the walls of electromagnetic coil 48, approximately halfwaythrough its motion along the length of hydraulic pump 42, andelectromagnetic coil 50 has been energized by current from battery 62.The continued motion of driving magnet 40 has placed further pressure,via hydraulic fluid 52, on annular driven magnet 44. Due to magneticinteractions with annular driven magnet 44, valve-seat magnet 54 hasmoved approximately halfway through its motion along the length of L-VAD74. Still closed, one-way valve 70 has pumped more blood, that wouldotherwise not have been pumped by the native heart, out of the leftventricle along the length of the ascending aorta 88 and into aorticarch 80. Aortic valve 76 remains open, allowing the flow of blood fromthe left ventricle 90 into ascending aorta 88.

[0116]FIG. 3 depicts the state of BEEP system 35 at the beginning of thereturn stroke of valve-seat magnet 54. As driving magnet 40 reverses itsprevious motion along the length of hydraulic pump 42, the flow ofhydraulic fluid 52 through main hydraulic line 60 is reversed as well,as indicated by the flow arrows in the Figure. The reverse flow ofhydraulic fluid 52 places pressure on annular driven magnet 44, pushingit back along the length of the L-VAD 74, in the direction of aorticvalve 76. As annular driven magnet 44 moves back along the length ofL-VAD 74, valve-seat magnet 54 and one-way heart valve 70 move in acorresponding manner. One-way valve 70 is open as it moves toward aorticvalve 76, allowing blood to flow freely through one-way valve 70 as itmoves. Aortic valve 76 is closed at this time, preventing blood fromflowing out of the L-VAD 74 and into left ventricle 90.

[0117]FIG. 4 depicts the state of BEEP system 35 halfway through thereturn stroke of valve-seat magnet 54. Driving magnet 40 has moved backwithin the walls of electromagnetic coil 48, approximately halfwaythrough its return motion along the length of hydraulic pump 42. Thecontinued-motion of driving magnet 40 has placed further pressure, viahydraulic fluid 52, on annular driven magnet 44, pushing it back downalong the length of L-VAD 74. Valve-seat magnet 54 has movedapproximately halfway through its return motion along the length ofL-VAD 74. One-way valve 70 is still open, allowing blood to flow freelythrough it as it moves. Aortic valve 76 remains closed, preventing theflow of blood from L-VAD 74 into left ventricle 90.

Pulsatile Flow and the Present Approach

[0118] The principles of fluid dynamics require a measurable work percycle (and power output) from the heart to overcome the pressuredifference in the passages of the circulatory system. Providingpulsatile instead of steady flow, accelerating and decelerating bloodand muscle, consumes significant measurable additional work (and power)from that required for steady flow. If the natural heart providedcontinuous flow under constant pressure, then thrombi would tend to formand gradually enlarge in relatively stagnant or low-velocity flowregions. In steady flow conditions these thrombi would tend to becomelarger with time. Eventually the larger thrombi could potentially bedislodged by the surrounding flow causing blockage in narrower passagesdownstream. The results would be disastrous. The human body would notprovide more pulsatile flow than that required for physiologicalreasons.

[0119] The human body requires pulsatile blood flow for survival, and asuccessful artificial heart pump or VAD should emulate the type ofpulsatile blood flow provided by the native heart. Unlike known artdevices, the present invention produces an optimized pulsatile flow. TheVAD of the present invention provides the “vector” or “matrix”difference between the unsteady flow required by the human body and theunsteady flow provided by the native diseased heart, hence supplyingonly the required deficit. By “vector” or “matrix” difference we implythat this is not a simple subtraction of two quantities, as it willbecome evident in the following. In a total replacement configuration(TAH) the invention provides the total unsteady flow required by thehuman body. While other inventions purport to optimize the flow, thepresent invention illustrates the actual requirements (engineeringprinciples) for this optimization.

[0120] The physical dimensions of the VAD or total replacement heartmust be optimized to each application (i.e. to each patient). The movingmass, damping and stiffness of the combined system (moving parts of theVAD plus native heart, if any, plus driven blood flow through thevessels plus hydraulic fluid, surrounding tissue, electromagneticdynamic phenomena, etc.) must be optimized to the dynamic response ofthe system (which is a form of the natural frequencies and damping ofthe overall system). If these conditions are not met, then the VAD ortotal replacement heart will be inefficient; it will require more powerthan the minimum to obtain the desired unsteady-flow output to the body.A good physical example of this is a yo-yo. If the string is pulled withthe right forces at the right times (which corresponds to the optimizedforcing function for the yo-yo), it requires minimum effort for maximumperiodic travel and produces spectacular results. If either the forcingfunction or the timing are not exactly right, then it takes more effortto obtain any travel, and the results are not as good. Another equallyimportant aspect of the invention is that the physical arrangement anddimensions of the invention are optimized to the desired amplitude andfrequency of the unsteadiness in blood flow required by the circulatorysystem.

[0121] Thus the power input to TAHs and VADs must be optimized to thedynamic response of the system, otherwise the efficiency will be low(they will require a lot of power to drive them). One of the claims ofthe proposed VAD is that its driving force-time and force-distancerelationships are optimized for minimum power input to the desiredunsteady-flow characteristics, via a prescribed procedure, thusincreasing its efficiency. This is done via a mathematical methoddescribed below. A pre-requisite for the use of this method is a deeperunderstanding of details of the flow and pressure conditions in thecardiovascular system than that in present medical and bioengineeringpractice. In other words, one needs to understand the details of thepressure and flow traces in the native as well as the artificial systemsin order to design an efficient VAD or TAH.

[0122] While it is understood that the pressure trace changes phase andamplitude downstream from the aorta, there is no acknowledgement as towhether the measured pressure traces are static, stagnation or totalpressures (defined in most fluids engineering texts). While it is clearthat during most of systole the left ventricular pressure must be higherthan the aortic pressure (otherwise flow would be in the reversedirection from the aorta to the ventricle), some texts indicateotherwise. The premise of this disclosure is that any TAH or VAD must beoptimized around the details of the pumping system and match therequirements of the human body.

[0123] Static pressure p_(st) is the pressure one would feel whiletraveling along with the velocity of the fluid in a channel. Stagnationpressure p₀ is the pressure one would feel with the fluid coming to restagainst the measuring device. Total pressure p_(T) is the stagnationpressure plus the static head of a column of fluid above the measuringpoint. For a perfect incompressible fluid of constant density ρ (whichis one of many frequently used mathematical models for blood) movingwith velocity C the governing equations are:

p ₀ =p _(st) +ρC ²/2

p _(T) =p _(st) +ρC ²/2+ρgz

[0124] The distinction between these three pressures in the blood flowis important in the design of the optimal VAD, as is the choice ofmeasurement devices that are specialized to distinguish measurement ofstatic, stagnation and total pressures, and the location of thesemeasuring devices in the system. Optimum design of the present device isintegrally related to the fundamental laws of fluid mechanics appliedfor unsteady flow conditions to the thermodynamic system enclosing theheart and circulatory system. Those skilled in the art of unsteadythermofluid dynamics will recognize that the system definition is ofparamount importance to the solution of the problem, and must be definedwith the accuracy and detail suggested in the text by Gyftopoulos andBeretta (1981); i.e. the system definition will require amounts andrange of valves for: matter, parameters or constraints, and interactingforces between system elements. These fundamental laws are usuallyexpressed as one equation for conservation of mass, three equations forconservation of momentum, for example, along (x,y,z), and a fifthequation for the energy balance (the first law of thermodynamics).

[0125] The following equations (1-5) are valid for any fluid continuum(compressible or incompressible, Newtonian or non-Newtonian), and theyare general in nature. The nomenclature used is as follows:

[0126] E_(t), e_(t)=energy and energy per unit volume,\\includinginternal, kinetic, and potential energy, etc.

[0127] e=specific internal energy

[0128] {right arrow over (f)}_(bd)=external body forcing function perunit volume (gravity, electromagnetic, etc.)

[0129] {right arrow over (f)}_(sf)=surface forcing function per unitvolume (resulting in stress tensor τ)

[0130] F_(nh){t}=force as a function of time from the native Heart

[0131] F_(vad){t}=force as a function of time from the VAD

[0132] h=specific enthalpy

[0133] m=mass

[0134] p=pressure

[0135] Q,q_(x),q_(y),q_(z)=heat into control volume, and heat per unitvolume in (x, y, z) coordinates

[0136] t=time

[0137] (x, y, z)=Cartesian coordinates

[0138] u, v, w=velocity components alone (x, y, z) coordinates

[0139] W=work into the control volume (from surface, from shaft, etc.)

[0140] x=axial direction

[0141] ρ=density

[0142] μ=dynamic viscosity

[0143] τ=stress tensor

[0144] τ_(ij)=element of stress tensor (includes pressure) along (i,j)coordinates

[0145] ∇=divergence operator

[0146] D=total derivative operator

[0147] ∂=partial derivative operator

[0148] The mass balance (continuity) is given by: $\begin{matrix}\begin{matrix}{{\frac{\partial\rho}{\partial t} + {\overset{\rightarrow}{\nabla}\left( {\rho \quad \overset{\rightarrow}{v}} \right)}} = 0} \\{{\frac{\partial\rho}{\partial t} + \frac{\partial\left( {\rho \quad u} \right)}{\partial x} + \frac{\partial\left( {\rho \quad v} \right)}{\partial y} + \frac{\partial\left( {\rho \quad w} \right)}{\partial z}} = 0}\end{matrix} & (1)\end{matrix}$

[0149] where the equation can be further simplified using certainassumptions such as incompressible fluid (but here we consider thegeneral form of the equation with no restrictions other than continuousfluid).

[0150] The vector form of the equation for conservation of linearmomentum can be written as the (x, y, z) momenta equations:$\begin{matrix}\begin{matrix}{{\frac{\partial\left( {\rho \overset{\rightarrow}{\quad v}} \right)}{\partial t} + {\overset{\rightarrow}{\nabla}\left( {\rho \quad \overset{\rightarrow}{v}\overset{\rightarrow}{v}} \right)}} = {{{\overset{\rightarrow}{f}}_{sf} + {\overset{\rightarrow}{f}}_{bd}} = {{\overset{\rightarrow}{\nabla}\overset{\overset{\rightarrow}{\rightarrow}}{\tau}} + {\overset{\rightarrow}{f}}_{b\quad d}}}} \\{{\frac{\partial\left( {\rho \quad u} \right)}{\partial t} + \frac{\partial\left( {\rho \quad u^{2}} \right)}{\partial x} + \frac{\partial\left( {\rho \quad u\quad v} \right)}{\partial y} + \frac{\partial\left( {\rho \quad u\quad w} \right)}{\partial z}} = {\frac{\partial\left( \tau_{xx} \right)}{\partial x} + \frac{\partial\left( \tau_{xy} \right)}{\partial y} + \frac{\partial\left( \tau_{xz} \right)}{\partial z} +}} \\{\quad {\overset{\rightarrow}{f}}_{x,{b\quad d}}}\end{matrix} & (2) \\{{\frac{\partial\left( {\rho \quad v} \right)}{\partial t} + \frac{\partial\left( {\rho \quad v\quad u} \right)}{\partial x} + \frac{\partial\left( {\rho \quad v^{2}} \right)}{\partial y} + \frac{\partial\left( {\rho \quad v\quad w} \right)}{\partial z}} = {\frac{\partial\left( \tau_{yx} \right)}{\partial x} + \frac{\partial\left( \tau_{yy} \right)}{\partial y} + \frac{\partial\left( \tau_{yz} \right)}{\partial z} + {\overset{\rightarrow}{f}}_{y,{b\quad d}}}} & (3) \\{{\frac{\partial\left( {\rho \quad w} \right)}{\partial t} + \frac{\partial\left( {\rho \quad w\quad u} \right)}{\partial x} + \frac{\partial\left( {\rho \quad w\quad v} \right)}{\partial y} + \frac{\partial\left( {\rho \quad w^{2}} \right)}{\partial z}} = {\frac{\partial\left( \tau_{zx} \right)}{\partial x} + \frac{\partial\left( \tau_{zy} \right)}{\partial y} + \frac{\partial\left( \tau_{zz} \right)}{\partial z} + {\overset{\rightarrow}{f}}_{z,{b\quad d}}}} & (4)\end{matrix}$

[0151] where the body forces are exerted on the whole body of fluid(such as by gravity, when {right arrow over (f)}_(bd)=p{right arrow over(g)}; or by external electromagnetic fields); and the surface forces areexterted by the interior surface of the control volume of fluid. Sometexts choose to separate the pressure terms from the stress tensor, buthere the pressure terms are included in the stress tensor τ.

[0152] The energy balance equation is given by: $\begin{matrix}\begin{matrix}{\frac{D\quad E_{t}}{D\quad t} = {\frac{D\quad Q}{D\quad t} + \frac{D\quad W}{D\quad t}}} \\\begin{matrix}{\begin{matrix}{\frac{\partial\left( {\rho \quad e_{t}} \right)}{\partial t} + \frac{\partial\left( {\rho \quad u\quad e_{t}} \right)}{\partial x} +} \\{\frac{\partial\left( {\rho \quad v\quad e_{t}} \right)}{\partial y} + \frac{\partial\left( {\rho \quad w\quad e_{t}} \right)}{\partial z}}\end{matrix} = {\frac{\partial\left( {q_{x} + {u\quad \tau_{xx}} + {v\quad \tau_{xy}} + {w\quad \tau_{xz}}} \right)}{\partial x} +}} \\{{\frac{\partial\left( {q_{y} + {u\quad \tau_{yx}} + {v\quad \tau_{yy}} + {w\quad \tau_{yz}}} \right)}{\partial y} +}} \\{\frac{\partial\left( {q_{z} + {u\quad \tau_{zx}} + {v\quad \tau_{zy}} + {w\quad \tau_{zz}}} \right)}{\partial z}}\end{matrix}\end{matrix} & (5)\end{matrix}$

[0153] where q_(x), q_(y), q_(z) are the external heat transfers (Q) ineach direction, the work (W) terms are given by tensor times velocityapplied to the surface of the control volume, and E_(t) includes all theenergy terms. For example, if these include only internal energy,kinetic energy, and potential energy, then E_(t)=ρe_(t)=ρ(e+|{rightarrow over (v)}|²/2+{right arrow over (g)}·{right arrow over (r)}).However, in the general case, E_(t) includes all energy terms affectingthe solution of the equations.

[0154] The above five equations can be written in vector form as:$\begin{matrix}{{\frac{\partial G}{\partial t} + \frac{\partial A}{\partial x} + \frac{\partial B}{\partial y} + \frac{\partial C}{\partial z}} = {\frac{\partial X}{\partial x} + \frac{\partial Y}{\partial y} + \frac{\partial Z}{\partial z}}} & (6)\end{matrix}$

[0155] where

G=[ρ, ρu, ρv, ρw, ρe_(t)]

A=[ρu, ρu², ρuv, ρuw, ρue_(t)]

B=[ρv, ρuv, ρv², ρwv, ρve_(t)]

C=[ρw, ρwu, ρwv, ρw², ρwe_(t)]

X=[0, τ_(xx) +{right arrow over (f)} _(x, bd), τ_(xy), τ_(xz) , q _(x)+uτ _(xx) +vτ _(xy) +wτ _(xz)]

Y=[0, τ_(yx)τ_(yy) +{right arrow over (f)} _(y, bd), τ_(xy) , q _(y) +uτ_(yx) +vτ _(yy) +wτ _(yz)]

Z=[0, τ_(zx), τ_(yx), τ_(zz) +{right arrow over (f)} _(z, bd) , q _(z)+uτ _(zx) +vτ _(zy) +wτ _(zz)]

[0156] (the native heart and VAD forcing function are terms {right arrowover (f)}_(x, bd) in X, Y, Z).

[0157] The only restriction in the above equations 1-6 is that bloodbehaves as a continuous fluid (if it didn't, e.g. if there iscavitation, severe lysis, or severe clotting, then the model isinadequate, but the resulting VAD is also useless). These equations arevalid for steady and unsteady flow (periodic and transient), with anyexternal force field or surface forcing function, with heat transfer,with Newtonian or non-Newtonian fluids etc.

[0158] The resulting instantaneous equations of fluid motion (1-6) haveinstantaneous eigenvalues and eigenvectors that can be computed, andthose must be matched with the combined forcing function from the nativeheart F_(nh{t}) and the VAD F_(nh{t}), i.e. the dynamic system ofequations for the native cardio-rheology as modified by the presence ofthe operating VAD. The resulting instantaneous system of dynamicequations are of the form:

[M]{{umlaut over (x)}}+[C]{{dot over (x)}}+[K]{x}=F{t}=F _(nh) {t}+F_(vad) {t}  (7)

[0159] where [M], [C], [K] are the instantaneous non-linear mass,damping and stiffness matrices respectively of the dynamic model. Theyare non-linear because they change with time and with mathematical orexperimental data model, because the human tissue and mechanicalcomponents are not linear, and because they change with instantaneousposition and geometry (for example with open or closed valves), and alsowith daily condition of the patient. In any case the procedures to modelthe dynamic system are well established, and the fidelity of the dynamicmodel is improving with time as better experimental data and theoreticalor numerical models become available for each component of the dynamicsystem.

[0160] The forcing function of the native heart F_(nh){t} to the dynamicsystem is provided by the human and can be measured (though it also canbe modeled with basic physiological interactions). The forcing functionof the VAD system is provided by the magnetic field to the coils, whichis generated by the current and voltage to the coils, so that for adiscretized dynamic system the instantaneous power (at any time t) bythe VAD is balanced with:

W(t)=F _(vad) {t}·{{dot over (x)}}+losses=V{t}·i{t}  (8)

[0161] where W(t) is the instantaneous power at any instant in time t,{{dot over (x)}} are the elemental velocities at the displacements whereelemental forces F_(vad){t} are acting, and the product V{t} i{t}represents the sum of the electric power (voltage times current)supplied to the coils. In one embodiment of the optimization procedurethe physical dynamic systems are linked so that the left-hand side ofequation (7) is linked directly in the optimization process to theright-hand side of equation (8). The losses are electromagnetic lossesof transmitting magnetic flux from the coils to the magnets, andfriction losses until this power reaches the elemental displacements {x}on which forces F_(vad){t} act, and other similar losses. These lossescan be measured or modeled mathematically with techniques available inmechanics, fluid dynamics, electromagnetism, and other engineeringtexts. Thus the model includes muscle, tissue, blood, hydraulic fluid,and electromagnetic and mechanical effects of mass, damping andstiffness. For example, these include friction and leakage in themechanical components and fluid passages, the hysteresis loop of theelectromagnetic drive of the VAD (a condition commonly called“latching”), other electromagnetic losses, and the stress tensors inequations 1-6, so that the resulting fluid-structure system is driven inan optimal manner.

[0162] In general engineering systems, the power optimization andcontrol-scheme optimization (such as those described later for themechanical blood pumping device and patient) would be best applied tothe actual system itself. In this particular system, namely the patientwith the mechanical pump surgically implanted, it would be extremelydifficult to perform the steady-state power optimization scheme, anddifficult to perform the control-optimization scheme, as this mayendanger the life of the patient. These best preferred embodiments ofusing the physical dynamic should eventually be possible after clinicaltrials. Alternative embodiments (alternative models) of the physical.dynamic system are likely to be used for scientific development of theBEEP system. These are likely to use analytic, numerical, orexperimental, etc., expressions, or their combination, to represent thephysical dynamic system. These models can be of various degrees ofcomplexity. Some of these models may represent the whole dynamic system,and others may represent only portions of the whole system. From theabove it is easy to foresee that one group of such possible dynamicmodels may include the mechanical blood pump only, while others may, inaddition, incorporate portions of the native heart and the circulatorysystem, etc. Similarly, one set of such models may concentrate onfinding the optimum F_(vad){t}, represented, for example, by forces andvelocities acting on one of: a) valve seat magnet 54; b) driven magnet44; or c)driving magnet 40, etc. With similar dynamic models of theelectromagnetic and hydraulic systems, this forcing function,F_(vad){t}, can be correlated with the instantaneous electric power tothe coils (resulting in a form of equation (8)). The preferredembodiment of the dynamic system model directly correlates the forces onthe left-hand side of equation (7) with the power on the right-hand sideof equation (8). Alternative embodiments of the optimization schemes mayuse simplified portions of the whole dynamic system, such as those thatfind the force, F_(vad){t} on: a) valve-seat magnet 54; b) driven magnet44; or c) driving magnet 40.

[0163] In this invention the forcing function F{t} in equation (7)consists of two parts, one provided by the native heart and the otherprovided by the VAD. For this purpose, “optimize” means minimizing thepower required to drive the VAD while minimizing the shear stressimposed on the blood.

[0164] Again, the terms “optimize” and “complement” are used inreference to the devices and systems of the present invention, it ismeant that at each heart beat and stroke of the VAD (used here to meaneither the VAD or TAH as described below), several actions are carefullytimed such that:

[0165] a) the native heart is allowed to pump as much blood as it can onits own before the VAD is activated;

[0166] b) as the blood-ejection phase of the native heart nearscompletion, the VAD is energized to provide additional pumping action;

[0167] c) the additional pumping action reduces the back pressure inthat native ventricle so that the native ventricle pumps more than itwould have pumped unaided;

[0168] d) the timing of the action, length of pumping stroke, and rateof pumping (stroke displacement versus time and resulting power inputversus time) of the VAD are related to the native heart ejected bloodvolume and rhythm in a manner that minimizes power input to the VADwhile meeting physiological constraints;

[0169] e) the optimization processes in d) take into account the dynamicinteraction between the native heart and the VAD; and

[0170] f) the optimization process and the control scheme are integratedwith the resulting changes in blood ejected per heart beat and heartrate (beats per minute) by the combined action of the native heart andthe VAD.

[0171] Specifically, the combination of the patient's nativecardiovascular system and the VAD at any condition of flow rate andbeating frequency supplied by the native heart will result in an optimalshape (function of location and time as shown in the Figure) for theforcing function provided by the VAD. The forcing function and frequencyof the VAD are controlled as explained elsewhere in these documents. Theequations presented above are general and they are not dependent on thedetails of the mathematical models. Some research teams will choose tosimplify these equations using the incompressible fluid approximation,Newtonian fluid approximation, linear models in finite element methodprograms or linearized equations in computational fluid dynamics (CFD)approaches. All of these simplifications are fully included in thegeneral equations (1-7).

[0172] The L-VAD is intended to be placed between the aortic root andthe aortic arch. Thus, for VAD applications the length L and overalloutside diameter D₀ of L-VAD 74 are limited by human physiology. Thereis a desire to directly wrap coils 46, 48 and 50 around the length oftravel of driven magnet 54, but this not always possible, due togeometric constraints. For example, for an adult male L≈10 cm and D₀≈4cm, the overall force that can be carried by a hollow magnet 54 is afunction of the volume of the magnet, among other factors, for example,if cylindrical, approximated by π(D₂ ²−D₁ ²)l/4, where D₂ and D₁ are theoutside and inside diameters of the magnet and l the length, themagnetic properties of the materials (factor k1), geometry (factor k2)and technology of components (e.g. leakage and friction characteristics,coil packing, heat transfer constraints (factor k3)).

f=F(k1,k2,k3)

[0173] where clearly, other factors being equal, the force is increasedby increasing D2, the outside magnet diameter. Thus the outside diameterof electromagnetic coils used in the linear magnet motor of prior artbecomes too big for VAD to fit into the human body in the vicinity ofthe aortic arch. This maximum-diameter issue is resolved with the use ofdriving magnet 44, the hydraulic fluid, and solid (or substantiallysolid) magnet 40. The following table is an indication for severaldistinct sizes of VADs, assuming that the diseased native heart provides50% of the cardiac output required by the human body: Normal Output ofRequired Cardiac Diseased VAD Weight Height Area Output Native Output(kg) (m) (m²) (cc) Heart (cc) (cc) Child 50 1.3 1.3 58 29 29 Teen 551.65 1.6 72 36 36 Avg. 55 1.75 1.7 76 38 38 Adult Female Avg. 75 1.852.0 90 45 45 Adult Male Large 110 175 2.3 102 51 51 Adult

[0174] As can be seen by the equations above, several standard sizes ofL-VAD 74 can be designed. As a guide, the smallest would be a pediatricdevice and the largest would be for a large adult. What follows is anexample of calculations performed on a hypothetical individual, and isnot intended to in any way limit the present invention.

[0175] The height and weight features of a person can be converted tobody surface area by using the approximate formula below. Once bodysurface area is known, normal cardiac output for a given individual canbe calculated. Normal cardiac output volume per body surface area is 45cc/m².

[0176] Body Surface Area (m²)=[ht (cm)]^(0.718)[wt(kg)]^(0.427)[0.007449]

[0177] Using a 75 kg 185 cm adult male as an example, the body surfacearea calculation results in a value of 2 m². Given the body surface areavalue of 2 m² calculated above, the normal stroke volume for theindividual is 90 cc. In end-stage cardiomyopathy, the native heartprovides approximately 50% of the required cardiac output. In theexample above the native heart would provide approximately 45 cc.Therefore, the L-VAD would have to provide an additional 45 cc.

[0178] What follows is a general description of the approximate sizes ofan L-VAD for the above patient at one ejection volume (45 cc) and oneheart rate. The procedure must be repeated several times for differentejection volumes and heart rates before the optimum L-VAD dimensions forthe patient are decided. The procedure is also affected by the size ofavailable one-way valves 70, especially if these are of the standardartificial heart valves available commercially, which are available inseveral standard diameters, usually measured in millimeters (mm).

[0179] The maximum L-VAD displacement required in this example is 45 cc.A standard 29 mm valve nomenclature is used for one-way valve 70. Thischoice affects the length of L-VAD 74 as well as the force that mustdrive driven magnet 44 and valve-seat magnet 54. The axial length ofdriven magnet 44 and valve-seat magnet 54 is 13 mm. The wall thicknessbetween driven magnet 44 and valve-seat magnet 54 is 1 mm, as is thethickness of outer sleeve 72.

[0180] For certain illustrative example cases the steady state andacceleration force required to pump blood through L-VAD 74 is 30 to 36 N(kg m sec⁻²). This is based on an initial estimate of the pressure thatwill be supplied by L-VAD 74, multiplied by the area of the pumpingdiameter. This initial estimate accounts for the acceleration of fluids(blood and hydraulic) pumped in the system (about 6 L in circulation),and the initial masses of the moving components. The volume of rareearth magnet in driven magnet 44 and valve-seat magnet 54 required toprovide the 30 to 36 N is 3.83×10³ mm³. The resulting thickness ofvalve-seat magnet 54 with a length of 13 mm is about 3.2 mm. Thus, theinside diameter of valve-seat magnet 54 is 29 mm and the outsidediameter is 35.4 mm. The inside diameter of driven magnet 44 is 37.4 mm.The thickness of driven magnet 44, with a length of 13 mm and an insidediameter of 37.4 mm, must be around 2.8 mm to achieve the desired 30-36N. Thus, the outside diameter of driven magnet 44 is 43 mm. Therefore,the outside diameter of L-VAD 74 is 45 mm.

[0181] The stroke length required if one-way valve is to give therequired 45 cc volume is 45.7 mm. Adding this to the axial length ofvalve-seat magnet (as required by the geometry of the device) theoverall axial length of the pumping portion of L-VAD 74 becomes 58.7 mm.This length will be increased to allow for the cuffs for hydraulic fluidand blood. It is understood that several alternative embodiments for thecross-sectional shape of the heart valve 70, magnets 54, 44 and 40, andcuff (or “capsule”) designs for the hydraulic connections for hydraulicfluid and blood can be used and will be apparent to one skilled in theart and thus they are hereby incorporated in this disclosure.

[0182] The above dimensions are used to provide geometric inputs for themodels used in equation (7). The inputs result in elements for massmatrix [M], damping matrix [C], and stiffness matrix [K]. Elements of[M] are evaluated using material densities. Elements of [C] areevaluated using fluid dynamics for the flow passages, structural dampingfor tissues and electromagnetic properties for magnets, coils and othercomponents, as needed. Elements of [K] are evaluated using material andsurrounding tissue properties and electromagnetic properties formagnets, coils and other components, as needed. The surrounding tissuemust extend to the control volume of the system where the tissuegeometry is not moving. This means a little further out of thepericardium (to fully include pericardium tremors) and a little furtherout of the blood and hydraulic fluid vessels (to include stiffness andcompliance, providing elements for [C] and for [K]).

[0183] For example, the pressure drops in hydraulic lines 60 and 82initially can be estimated using analytic calculations available instandard textbooks, and later evaluated by discretized mathematicalmodels as elements of matrices in equation (7).

[0184] Continuing the above example, with certain engineeringassumptions driving magnet 40 could be 3800 mm³ grade 37 rare earthmagnet. In one embodiment this magnet could have radius 9.7 mm, andlength 12.86 mm. The hydraulic volume displaced by 45.7 mm of strokelength of driven magnet 44 is 54.7 mm. Thus, the overall length ofhydraulic pump 42 is about 81 mm (this length will be increased by thelength of the hydraulic cuffs).

[0185] Additional secondary calculations are made to evaluate thegeometries of auxiliary components such as hydraulic lines, and othercomponents such as tissue in the myocardium and blood flow system. Theinputs result in elements for matrices [M], [C], and [K] (some measuredin clinical trials, others measured for individual patients). Theelements of vector of displacements {x} and its derivatives {{dot over(x)}} and {{umlaut over (x)}} in equation (7) are elementaldisplacements. The equation is nonlinear and can be decomposed in a fewor infinitely many degrees of freedom, depending on the fidelity of thedynamic model.

[0186] F_(nh){t} in equation (7) is measured for each condition (heartrate, ECG signal, volumes ejected from right and left ventricles, andpressures) of the patient. For L-VAD 74 this is given by the totalpressure (static+dynamic+elevation components) provided by the diseasedheart inside the left ventricle as a function of time (measured duringthe heartbeat) integrated over the inside surface area of the fourchambers of the heart. This surface area is also measured with magneticresonance imaging (MRI), echocardiography, or other similar technique.These give pressure-volume-time traces for the diseased heart asillustrated in FIGS. 14 and 15. The volume information is correlatedwith data from proximity sensors, such as 406 and 416 in FIG. 9, whichmay be, for example, proximity sensors. This information changes as thecondition of the patient worsens or improves. This means that the dataneeds to be calibrated before surgery, and again soon after surgery, andmonitored periodically so that the data provided by proximity sensors406 and 416 reflect the forcing function provided by the native heart,where the mathematical expressions are:

dF _(nh) {t}=p(t)dA(t)

F _(nh) {t}=∫ _(A) p(t)dA(t)

[0187] The resulting pressure-volume-time traces of the native hearthave a phase associated with the timing of the forcing functionF_(nh){t} during the beat. This can be modeled with Fourier seriesanalysis of the pressure and volume signals of the native heart overtime. Again, these vary with the rate (beats per minute) of the nativeheart and with the condition of the patient (i.e. the informationchanges as a function of time and needs periodic updating).

[0188]FIG. 6 is a cross-sectional view of L-VAD 74, along line 6-6 ofFIG. 1. By contrast, FIG. 7 is a concept illustration of the human heartshowing the location of L-VAD 74 in place of ascending aorta 88. It isunderstood that the L-Vad may replace a portion, not necessarily theentire ascending aorta. Further in some embodiments the aorta may simplybe transected to place the L-VAD outside the body, with blood conduitsconnecting the ends of the transected aorta to the L-VAD. Correspondingalternative embodiments are possible for the R-VAD, BI-VAd. Blood movesfrom left ventricle 90, through aortic valve 76 and into L-VAD 74,situated in place of the ascending aorta 88, pumps blood into aorticarch 80.

[0189]FIG. 8 shows the placement of an entire BEEP system 35 within thehuman torso. The illustration depicts the spatial relationship betweenbattery/controller assembly 65 and L-VAD 74. FIGS. 7 to 13 and 20 to 25are schematic illustrations, not cross-sectional views, and the locationof L-VAD 74 in FIG. 8 is at a different plane from the location of R-VAD58 in FIG. 21. (The L-VAD in FIG. 8 is correctly shown more to the rightof the patient's chest than RVAD in FIG. 21, but FIGS. 8 and 21 areanatomically correct, while FIGS. 7 and 20 are simple arrangementillustrations).

[0190]FIG. 9 is a concept illustration of the human heart, illustratinga technique to measure the volumes of the left and right ventricles,which is used in the control algorithm. Rare earth (or similar material)magnets 402 and 404 are embedded in the endocardial surface of the rightventricle, and their relative motion changes the magnetic field betweenthem. These changes are measured by proximity sensor 406, mounted on thepericardium. The signal is transferred by electrical lead 408 to wirebundle 410. Rare earth or similar material magnets 412 and 414 areembedded in the endocardial surface of the left ventricle, and theirrelative motion changes the magnetic field between them. These changesare measured by proximity sensor 416, mounted on the pericardium. Thesignal is transferred by electrical lead 418 to wire bundle 410. Thesignals from wire bundle 410 are transmitted to controller 64 and usedas described later.

[0191] The proximity sensors are currently available devices that mayoperate on the resistive, capacitative or inductive principles, orcombinations, or other similar distance-measuring technology. Auxiliary(parallel horizontal) lines 1 through 6 in FIGS. 10 through 13 representthe motion of the ventricles and the KG diaphragm (see below) during acardiac cycle. It has traditionally been thought that the valves of theheart open to let the blood through when the chambers contract, and snapshut to prevent it from flowing backward as the chambers relax. Whilethis is correct, the valves also act as pumping pistons for at least aportion of the cardiac cycle, a fact not known to be previouslyrecognized in the literature. The plane of the valves and the supportingtissue on the perimeter of the valves form an internal diaphragm,approximately in the horizontal plane, which buckles and moves in 3D,which is also not known to be named in the existing literature. For thepurposes of this document this diaphragm will be theKorakianitis-Grandia (KG) diaphragm, illustrated in FIGS. 10 to 13, andgenerally designated at 92. KG diaphragm 92 has four quadrants with avalve in each quadrant. It is activated by the surrounding cardiacmuscle, which forces the diaphragm into a periodically-changingthree-dimensional surface. (Thus FIGS. 10-13 are illustrations ratherthan cross-sections of the human heart). During the cardiac cycle theaortic valve and pulmonary valve stay nearly immobile (which allows oneto place the VAD on the outlet side of these two valves); but the mitraland tricuspid valves move substantially, contributing at least in partto the pumping action of the ventricles. The mitral valve movement iscomparable to the movement of the inside wall of the left ventricle. Thetricuspid valve exhibits an even greater excursion and correspondingpumping action and is actually used in current medical practice as ameasure of right ventricular ejection fraction.

[0192] While the exterior surface of the heart moves slightly during thecardiac cycle, the volume of the four-chamber heart does not changeappreciably in time. However, the known art does not recognize that thetotal overall volume of each of the two sides, left and right, of theheart does not change appreciably during the cardiac cycle, even thoughthe ventricular and atrial septa move. In operation of the native heart,during left ventricular ejection, the left atrium concurrently expands(while filling for the next cycle), and KG diaphragm 92 begins to movetowards the apex of the heart, with complementing motions of the atrialseptum and of the ventricular septum, thus keeping the overall volume ofthe left side of the heart about constant. Apex of the heart is a commonterm for the tip of the left ventricle. Similar arguments keep theright-side volume approximately constant, while the right and left sidesof the heart expel blood to the lungs and the aorta aboutsimultaneously. Minor deviations from these equal-volume considerationson each side, right and left, occur due to one side of the heart beatingslightly before the other, heart-muscle and blood-vessel elasticity,transient accelerations or decelerations of the overall cardiac cycle(governed by the body's demand), and blood compressibility (which atcirculatory system pressures is practically negligible).

[0193] The human cardiac cycle consists of two phases, conventionallycalled diastole and systole. During diastole (FIGS. 10 and 13) theventricular muscle is relaxing, KG diaphragm 92 moves toward the base ofthe heart while the aortic and pulmonary valves are closed, and themitral and tricuspid valves are open and moving towards the base of theheart, thus increasing the volume inside the ventricles whileconcurrently decreasing the volumes inside the atria. Base of the heartis a common term for the posterior aspect of the heart, behind the atriain the heart's anatomical position. The open mitral and tricuspid valvesmove upwardly (when the body is upright) to engulf blood from what wasvolume inside the atria (thus concurrently increasing the volume insidethe ventricles while decreasing the volumes inside the atria). In amodel of ventricular flow this volume exchange would affect thethermodynamic system definition, mentioned previously. A scrutinizingreview of echocardiography tapes reveals that radial volume changesaround the vertical axes of the ventricles account for roughly 75% ofthe volume change, with the corresponding movement of the KG diaphragmaccounting for the remaining 25% volume change. There is heart-musclework associated with these changes in volume that must be accounted withthe correct mathematical model in any attempt to model flow in the heartor in VAD mimicking the function of the heart.

[0194] By the end of diastole, the relaxing ventricle allows KGdiaphragm 92 to move toward the base of the heart. During initialsystole (FIG. 11) the aortic and pulmonary valves remain closed whilepressure is building up inside the ventricles. Subsequently the bloodpressure inside the left ventricle becomes higher than the pressure inthe ascending aorta, and the aortic valve is opened by blood flowing outof the ventricle. Substantially concurrently the blood pressure insidethe right ventricle becomes higher than the pressure in the pulmonarytrunk, and the pulmonary valve opens. The motion of the KG diaphragmcarrying the mitral valve toward the apex of the heart, along with thesimultaneous concentric contraction of the ventricle, ejects blood intothe ascending aorta.

[0195] Correspondingly, the movement of KG diaphragm 92 carrying thetricuspid valve towards the apex of the heart, along with thesimultaneous contraction of the right ventricle, ejects blood into thepulmonary trunk. This same motion of the KG diaphragm with the tricuspidand mitral valves closed increases the volumes inside the atria, hencerefilling the atria with blood from the pulmonary veins (left atrium)and the vena cava (right atrium). Towards the end of systole the aorticand pulmonic valves close, then the mitral and tricuspid valves open andthe cycle starts anew. Each cycle takes approximately one second.

[0196] The double throb (“lub dub”) of the beating heart is generated bythe snapping of the closing valves, but also from the accompanyingvibrations of the surrounding heart muscle and contained blood.

[0197] The three-dimensional motion of KG diaphragm 92 forces each oneof the four one-way valves to act as pumping pistons for at least partof the cardiac cycle. The blood flow must be optimized around artificialheart valves to provide the desirable flow and pressure pattern whileminimizing shear stresses on blood cells.

[0198]FIG. 10 represents late diastole. At this point KG diaphragm 92 isat its uppermost position (horizontal lines 1-3). Pulmonary valve 94 andaortic valve 76 are closed, and mitral valve 96 and tricuspid valve 98are open, completing filling of ventricles 106 and 90 followingcontraction of right atrium 100 and left atrium 102. Ventricularmyocardium 104 is in its relaxed state.

[0199]FIG. 11 represents early systole. At this point ventricularmyocardium 104 is thickened concentrically, and KG diaphragm 92 ismoving downward (horizontal lines 2-4), the tricuspid valve side more sothan the mitral valve side. Due to this motion, the volume of ventricles106 and 90 is decreased while that of atria 100 and 102 is increased.Hence, the total volume of the heart remains essentially constant.

[0200]FIG. 12 represents late systole. At this point ventricles 106 and90 have maximally thickened concentrically, and KG diaphragm 92 has beenpulled maximally downward (lines 4-6). This completes the emptying ofthe ventricles.

[0201]FIG. 13 represents early systole. At this point KG diaphragm 92 isbeginning to return upward (lines 3-5) toward atria 100 and 102, whileventricular myocardium 104 relaxes concentrically. As a result of thismotion, the volume of atria 100 and 102 decreases while that ofventricles 106 and 90 increases, hence the overall volume of the heartremains essentially constant.

[0202] There are minor variations to the basic steps outlined above, dueto damping and elasticity of the heart tissues, and small amounts ofnative heart valve leakage, which can be accounted for in thethermodynamic system definition mentioned earlier.

[0203]FIG. 14 represents pressure-volume loops for healthy (solid lines)and diseased (broken lines) hearts, with pressure plotted along the Yaxis and volume plotted along the X axis. Line 112 represents a healthyheart. Point 118 to point 120 represents ventricular filling. Point 120to point 122 represents “isovolumetric contraction”. Point 122 to point124 represents ejection during systole. Point 124 to point 118represents isovolumetric relaxation.

[0204] Line 114 represents the normal response in accordance with theFrank Starling law to an increase in volume. Points 126, 128, 130 and132 correspond to points 118, 120, 122, and 124, on line 112,respectively, and represent the corresponding phases of the cardiaccycle. Notice that points 124 and 132 lie on the same line, commonlyreferred to as the End Systolic Pressure Volume Relationship (ESPVR).

[0205] Line 116 represents a diseased heart which has exceeded thelimits of the Frank Starling curve. In these hearts, the end diastolicpressure and volume are elevated but the end systolic pressure isdecreased from that associated with normal myocardium, as noted by alesser slope of the ESPVR line. Points 134, 136, 138 and 140-correspondwith points 118, 120, 122 and 124 of line 112, respectively.

[0206]FIG. 15 represents ventricular pressure over time of the healthyand diseased hearts. Again, the end diastolic ventricular pressure isgreater in the diseased heart than in the healthy heart, and because theejection fraction is decreased in the diseased heart, the heart rate isincreased so that the total cardiac output is maintained.

[0207] VADs are activated by either ECG signal, or via a fill-to-emptymode. The control algorithm of the present device utilizes inputs fromthe ECG of the native heart, as well as a measurement of the ejectionvolume and pressures of both the right and left ventricles. As a result,the prior art complication of mismatch of ejection volume between theright and left ventricles is eliminated.

[0208]FIG. 16 is a concept illustration of the pumping travel of newBEEP system 35, as compared to that of known cardiac pumping devices.The abscissa is the number of beats per minute of the native heart. Theordinate is the length of travel of driving magnet 40 along the lengthof hydraulic pump 42. The known devices are either on or off (travel Xis always equal to maximum travel Xmax), as shown by the dotted line. Inthe present device, the length of travel of driving magnet 40 alonghydraulic pump 42 varies, following solid line 142, depending on thenumber of beats per minute of the native heart. As the number of beatsper minute of the native heart increase, they reach high thresholddotted line 144 at which point controller 64 signals for driving magnet40 to start moving along the length of hydraulic pump 42 smoothlyincreasing the stroke travel, approaching solid line 142.

[0209] Over time, due to augmentation of ejected volume by the VAD, theend diastolic volume of the native heart decreases. This allows theinternal volume of the ventricle to become smaller, which subsequentlyallows the muscle of the native heart to begin to recover, and henceeject a greater volume of blood per stroke. As cardiac output is theproduct of ejected blood volume times heart rate, the increased ejectedvolume allows the heart rate to decrease. These changes are sensed bycontroller 64, which reduces the stroke length of the VAD as a greaterportion of the cardiac output is now being supplied by the native heart.The stroke length of the VAD progresses to the left, from point 148toward point 146. When the beats per minute reach the lower thresholdpoint 146 (reflecting at least partial recovery of the nativemyocardium), this will cause a decrease in stroke length along line 150,which effectively reduces the stroke length of the VAD. If recovery ofthe native heart continues, the stroke volume of the L-VAD is reduced tozero along line 150. At this point the native heart is once againproviding the total cardiac output on its own, without the assistance ofthe VAD. The shape of line 142 is actively manipulated by controller 64using additional inputs for the measurement of the ejection volume andpressures of the right and left ventricles provided by the measurementsystem shown in FIG. 9 and described in relation thereto.

[0210] A critical factor in the success of a newly-installed L-VAD issatisfactory operation of the right ventricle during the immediateperi-operative period. The ejection fractions of the right and leftventricles are monitored by mechanisms such as those illustrated in FIG.9 and alternative embodiments thereof. This allows the volumetricoutputs of the right ventricle and the assisted left ventricle to be thesame by manipulating the shape of line 142 up or down. For example,suppose that a short time after activation, the right ventricle ismeasured to give 50 cc per beat and the left ventricle gives 25 cc perbeat. In this case, the stroke length of L-VAD 74 will be adjusted togive 25 cc per beat. If, a short time later, the right ventricle startsto fail and now only ejects 40 cc per beat (while the left ventriclestill gives 25 cc per beat), this discrepancy in the ejection fractionswill be detected by proximity sensors 406 and 416. In response,controller 64 will lower the level of line 142, resulting in a shorterstroke length of L-VAD 74, to give 15 cc per beat, for a total of 40 ccfrom the assisted left ventricle. The human body will compensate with acorresponding increase in heart rate (beats per minute). In thisfashion, controller 64 matches the ejection volumes of the right andleft sides. Should the right ventricle continue to fail, a decision willhave to be made as to the appropriateness of installing an R-VAD, makingthis a BI-VAD system.

[0211] Assuming that a patient is supported on an L-VAD alone, thelength of the pumping stroke of the L-VAD is determined as a function ofbeats per minute, as shown in FIG. 16, and manipulated by matching theejection volumes of the left and right sides of the heart. For example,starting from point 148, if the beats per minute continue to increase,then the piston stroke also continues to increase smoothly to point 178.Starting from any operating point to the right of point 178, as beatsper minute decrease, at point 178 the device reduces the travel ofdriving magnet 40 from its maximum travel along hydraulic pump 42. Ifthe number of beats of the native heart is reduced sufficiently to reachlow threshold line 150, then the travel of driving magnet 40 is reducedsmoothly until it becomes zero following low threshold line 150. Lines150 and 142 may coincide over the length of line 150. Line 142 to theleft of point 146 in FIG. 16 can represent the initial activation of theL-VAD after surgical installation thereof.

[0212] The locations, magnitudes and exact shape of lines 142, 150 and144 shown in FIG. 16 are for purposes of illustration and will vary frompatient to patient, and device to device. In addition, during normaloperation of the L-VAD, small up or down variations of the level of line142 are made by controller 64 in order to match the ejection volumes,measured as described in FIG. 9, of the left and right sides of thesystem. The control algorithm has several input variables; among others,beats per minute measured by the ECG (as described below), the ejectionvolume and pressures from the right and left side of the -system asillustrated in FIG. 9. For clarity, in the remaining Figures theconcepts are illustrated using beats per minute to represent thefunction of controller 64.

[0213]FIG. 17 is a chart comparing the activation sequence of the knownLarson et al. device and that of the present BEEP system, as well as thecorresponding power requirements, in relation to the ECG trace of thenative heart. The solid lines represent the present device and thedotted liens represent the known device of Larson et al. The commonabscissa is the time period required for two beats of the native heart.The ECG trace has characteristic spikes Q, R, S (commonly known as theQRS complex), and waves T and P, whose physiological function andimportance is described in detail in medical texts. The beginning andend of the stroke of Larson's device occurs at or near point R of theQRS signal.

[0214] The graph shown in portion A of FIG. 17 shows non-dimensionalstroke distance traveled by driving magnet 40 (X to Xmax) from 0.0 to1.0, according to FIG. 16. Portion B of FIG. 17 illustrates a typicalECG voltage trace during cardiac operation. The beginning of the strokeof the present system is at or about the beginning of the T wave,allowing the rapid ejection phase of the native heart to precede theaugmentation of the VAD. The pumping phase (points 154 to 156) of thepresent system occurs between the end of the T wave and the beginning ofthe P wave. The return stroke begins at this time (point 156) and endsat or just after the QRS complex (point 158).

[0215] Driving magnet 40 rests at the center of electromagnetic coil 46(X=0) during the time period between the end of the return stroke (point158) and the beginning of a new stroke (point 160). The resting periodbetween the end of the return stroke (point 158) and the beginning ofthe next stroke (point 160 corresponding to 154) is important for anumber of reasons. For example, acceleration at the beginning 154 andend 156 points of the stroke is minimal. This resting period allows timefor depolarization of electromagnetic coils 46, 48 and 50 betweenstrokes. When necessary, it also allows driving magnet 40 to be centeredwithin electromagnetic coil 46, thereby allowing driven magnet 44 andvalve seat magnet 54 to return to the beginning of the stroke of L-VAD74. This return function is accomplished by opening and closing checkvalve 84, as necessary.

[0216] Due to leakage of hydraulic fluid around magnets 40 and 44 it ispossible that one of the two magnets is stopped at one of the two endsof its travel while the other magnet is somewhere in the middle of itsstroke. For example, if annular driven magnet 44 is at the end of itstravel at the pump inlet (by the aortic valve as shown in FIG. 1) butdriving magnet 40 is not yet all the way back to the beginning of itsstroke (as shown in FIG. 4), then high hydraulic pressure will arisebetween driving magnet 40 and end cap 57. This condition would be sensedby the large increase in the power required by the coils. At that timethe controller would open check valve 84 and would pull driving magnet40 by the coils towards end cap 57, until it touches end cap 57, thusbringing the two magnets back into phase, and normal operation wouldresume. The procedure is similar if driving magnet 40 reaches the end ofits pumping travel (as shown in FIG. 3) while driven magnet 44 is nearthe middle of its travel (as shown in FIG. 2). It is also possible tocorrect for these leakages at the end of every pumping stroke, or everfew pumping strokes. A similar procedure can be used for initialactivation of the device, to start the device after it has been stopped,and to re-lock magnets 54 and 44 if they are not locked relative to eachother at any time during operation. The latter condition is sensed by alarge decrease in the power required by the coils.

[0217] The shape of line 164 in FIG. 17 (portion A) is determined by theoptimization procedure described later herein. Acceleration beginssmoothly (point 154 on line 164) so that less power is required than ifthe device started with a constant velocity. Maximum acceleration isachieved somewhere in the middle of the stroke, based on theoptimization procedure. The, VAD of the present system approachesmaximum stroke travel with minimum velocity at point 156 so that it doesnot impact against the mechanical stop at the end of travel and noenergy is lost due to impact. Thus, less energy is required to start thereturn stroke. The return stroke is less critical than the pumpingstroke because one-way valve 70 is open and less energy is required toreturn the new VAD (e.g. L-VAD 74) to its starting position. Even so,the shape of line 162 is optimized by the procedure. Velocity is zero atX=Xmax, requiring less power than if there was a change in velocity atXmax. The shape of lines 164 and 162, and resting period between points158 and 160 is optimized by the procedure.

[0218]FIG. 17C shows the power requirements of the present device inWatts during usage. The maximum power requirement (point 166) occurssomewhere along line 164, as optimized by the procedure. During thereturn stroke, power peak at point 168 occurs slightly before X=Xmax atpoint 156, which corresponds to the power requirement at point 170. Thisoccurs because of the sequence of energization of electromagnetic coils46, 48, and 50 as explained further in FIG. 18.

[0219] In comparison, the known device, illustrated by the dotted linesin FIG. 17, begins the pumping stroke at or near R of the QRS complex,with constant velocity, until the point of maximum travel, which occursat or near the end of the T wave. The return stroke is with constantvelocity from the point of maximum travel until the R peak of the nextQRS complex, requiring large acceleration at the two ends 172 and 174 ofits stroke. This requires correspondingly large power input. Inaddition, the velocities are not optimized for the unsteady flow and thetime varying magnetic fluxes, requiring large power input at all pointsduring the stroke. Points 172 and 174 correspond to power peaks 176 and178, respectively. As a result, the present device will take less power(solid line 180) than the prior art device (dotted line 182), and thepeaks occur at different times.

[0220] Further with reference to FIG. 17, power peak 166 of the presentdevice occurs during the T wave of the native heart, allowing the nativeheart to finish its rapid ejection. This increases the volume of bloodpumped due to the combination of the native heart and the VAD withrespect to the prior art (due to summation of volume), requires lesspower than the prior art device, and allows the native heart a chance torecover by decreasing left ventricular volume. These combinations makethe BEEP system bio-compatible.

[0221]FIG. 18 is a schematic illustration of the embodiment of BEEPsystem 35 shown in FIG. 1, using three electromagnetic coils 46, 48 and50 along the length of hydraulic pump 42 showing the correspondingmagnetic flux of the driving magnet and the electromagnetic coils.Vector 184 represents the magnetic flux of driving magnet 40. PositionX=0 is at the center of electromagnetic coil 46. Position X/Xmax is atthe center of electromagnetic coil 50. The top portion of the Figureshows electromagnetic coils 46, 48 and 50 and positive stops 186 and188. The middle portion of the Figure is an illustration of the typicalperiodic representation of the magnetic fluxes of electromagnetic coils46, 48 and 50 during the pumping stroke from X=0 to X=Xmax. The bottomportion of the Figure is an illustration of the typical periodicrepresentation of the magnetic fluxes of electromagnetic coils 46, 48and 50 during the return stroke from X=Xmax to X=0.

[0222] With reference to the middle portion of FIG. 18, the abscissa isthe axial position of driving magnet 40 from the center of coil 46(point 190) to the center of electromagnetic coil 50 at X=Xmax (point192).

[0223] With reference to the bottom portion of FIG. 18, the abscissa isaxial position of driving magnet 40 from the center of coil 50 at X=Xmax(point 192) to the center of electromagnetic coil 46 at X=0 (point 190).Starting from X=0 in FIG. 18(b) (point 190, which corresponds to point154 in FIG. 17), positive stop 186 ensures that activation of magneticfields 194 and 196 in electromagnetic coils 46 and 48, respectively,will force driving magnet 40 from the center of electromagnetic coil 46towards the center of electromagnetic coil 48. Magnetic field 194 isreduced to zero soon after driving magnet 40 is a little outsideelectromagnetic coil 46. As driving magnet 40 approaches the center ofcoil 48, magnetic field 196 in coil 48 is reversed in direction tomagnetic field 198. The reversal in magnetic field 196 does notnecessarily coincide with the point in time when magnet 40 is at thecenter of coil 48. The exact location of reversal is dependent upon anoptimization procedure.

[0224] Magnetic field 200 is initiated just before driving magnet 40enters electromagnetic coil 50. The combination of magnetic fields 198and 200 die out by point 192 (corresponding to point 156 in FIG. 17) andsmoothly bring the magnet to position X=Xmax, against positive stop 188.At that position, the magnetic fields in coils 48 and 50 are reversed asshown at point 192 in FIG. 18(c). Positive stop 188 ensures thatmagnetic fields 202 and 204 from coils 50 and 48, respectively, pushdriving magnet 40 from X=Xmax (point 192) towards the center of coil 48.Magnetic field 202 is reduced to zero soon after driving magnet 40 is alittle outside electromagnetic coil 50. As driving magnet 40 approachesthe center of coil 48, magnetic field 204 in coil 48 is reversed indirection to magnetic field 206. The reversal in magnetic field 204 doesnot necessarily coincide with the center of coil 48.

[0225] The exact location of reversal is dependent upon the optimizationprocedure. Magnetic field 208 is initiated just before driving magnet 40enters electromagnetic coil 46. The combination of magnetic fields 206and 208 die out by point 190 (corresponding to point 158 in FIG. 17) andsmoothly bring the magnet to position X=0, against positive stop 186. Atthat position, from point 158 to point 160 in FIG. 17, magnetic field208 (or its residual effects) will retain driving magnet 40 at X=0,whereupon the cycle repeats itself. All of the magnetic fields 194-208will be optimized to give F_(vad){t}. The power to obtain magneticfields 196-208 will be minimized based on the resistance, inductance,and capacitance of the electromagnetic system, including the coils andmagnets, and voltage source, using constitutive relations orexperimental data for the dynamic representation of the systems inequation (7), established in electromagnetic theory.

[0226] In general, the magnitudes of magnetic fields 194-200 will begreater than those of magnetic fields 202-208, because the former occurduring the pumping phase with one-way valve 70 closed and pushing blood,while the latter occur during the return stroke with one-way valve 70open. Even though the above embodiment utilizes three electromagneticcoils, it is contemplated that the present device may contain more orfewer electromagnetic coils. In the limit, a linear stepper motor may beused.

[0227]FIG. 19 is a schematic illustration of the embodiment of BeepSystem 35 illustrated in FIG. 1, wherein only two electromagnetic coils46 and 48 are activated and used to move driving magnet 40. This is donein order to obtain some stroke length X less than Xmax, as shown in FIG.16. Similar relative functioning, as described with respect to FIG. 18,is found in the displacements and magnetic fluxes of FIG. 19 as well.

[0228]FIG. 20 is a concept illustration of the human heart depicting theplacement of R-VAD 58 in place of a portion of the pulmonary trunk 210.Blood moves from right ventricle 106 through pulmonary valve 94 and intothe pulmonary trunk 210 before its bifurcation point 212.

[0229]FIG. 21 shows the placement of an R-VAD embodiment of BEEP system35 within the human torso O. The illustration depicts the spatialrelationship between battery battery/controller assembly 65, hydraulicpump 42, and R-VAD 58. As mentioned previously, FIGS. 7, 9-13 and 20,22, 24 and 27 are simple arrangement illustrations, notanatomically-correct views.

[0230]FIG. 22 shows the placement of a BI-VAD, generally designated 77,which consists of a combined assembly of L-VAD 74 and R-VAD 58 in thesame system. In this embodiment L-VAD 74 is located in place of at leastpart of the ascending aorta 88. In use of BI-VAD 77 blood moves fromleft ventricle 90 through the aortic valve 76 and into ascending aorta88; i.e. in this system L-VAD 74 pumps blood into the aortic arch 80,just as in use of the L-VAD alone. RVAD 58, as part of the BI-VAD 77, islocated in place of at least part of the pulmonary trunk 210, just as itis used in the embodiment (R-VAD alone) shown in FIG. 20. Blood movesfrom the right ventricle 106 through pulmonary valve 94 and intopulmonary trunk 210 as R-VAD 58 portion of BI-VAD 77 pumps blood intothe bifurcation 212 (hidden from view) of the pulmonary arteries.

[0231]FIG. 23 shows the general placement of the BI-VAD 77 embodiment ofthe BEEP system in the human torso O. The illustration depicts therelative spatial relationship of BI-VAD 77 in the chest and ofbattery/controller assembly 65 and hydraulic pump 42 in the abdomen.(The L-VAD in FIG. 23 is correctly shown more to the right of thepatient's chest than RVAD, but FIG. 23 is anatomically correct, whileFIG. 22 is a simple arrangement illustration).

[0232]FIG. 24 is a schematic illustration of the Total Artificial Heart(TAH) embodiment, generally designated 95, for use in a variation of thenew BEEP system 35. In this TAH embodiment, atria 100 and 102, alongwith ECG signal 66, are retained from the native heart. The TAH iscomprised of a BI-VAD system with a greater stroke volume than theassist embodiment, as the total cardiac output is now being supplied bythe BEEP system. In some TAH embodiments it is possible to use themitral valve 96 and tricuspid valve 98, shown in previous Figures,provided their papillary muscles and chordae tendineae are functioning,and insert he L-VAD and/or R-VAD portions of the BI-VAD into therespective ventricles. However, in other alternative TAH embodimentsartificial valves may be necessary or preferred. In addition, all, someor non of the native ventricle may be retained. In the embodimentillustrated in FIG. 24 L-VAD 74 completely takes the place of leftventricle 90 (seen in FIG. 7, for example), and hence its inlet isgrafted to an artificial valve (not shown) and its outlet is graftedinto the ascending aorta 88. Right ventricle 106 is replaced with anR-VAD 58, which has its inlet grafted to an artificial heart valve (notshown) and its outlet grafted into the pulmonary trunk 210.

[0233]FIG. 25 shows the general placement of the TAH embodiment 95 ofthe new BEEP system within a human torso O. The illustration depicts thespatial relationship among battery 62, controller 64, in the abdomen,and TAH 95, in the chest cavity. In various embodiments of the TAH aportion or all of the native heart may be removed. If the chordaetendinae and papillary muscles are intact, then the TAH would consist ofan L-VAD and an RVAD placed inside the respective ventricles. If themitral and tricuspid valves of the native heart are not utilized, thenthe TAH would require a one-way valve at the inlet of the L-VAD and aone-way valve at the inlet of the RVAD, and the combination that wouldmake the overall TAH. If the entire native heart (including thesino-atrial node) is removed, then the L-VAD and RVAD system that wouldcomprise the TAH will be triggered by an electrical signal driven bysensors that indicate the level of oxygen in the blood stream and othersensors of body functions.

[0234]FIG. 26 represents an alternative embodiment, generally designated35′, of an L-VAD version of BEEP system 35. In this version, two or moreelectromagnetic coils, 300, 302 and 304 are used to drive valve seatmagnet 54 in a reciprocal fashion. In this alternative embodiment, nohydraulic pump is required. The Figure represents the beginning of thepumping stroke, at which time one-way valve 70 is closed andelectromagnetic coils 300, 302 and 304 are energized in a manner similarto that illustrated in FIGS. 18b or 19 b to drive valve seat magnet 54down the length of L-VAD 306, pumping blood into aortic arch 80. At theend of the pumping phase, the direction of current in electromagneticcoils 300, 302 and 304 is changed, again in a manner similar to thatdepicted in FIGS. 18c or 19 c, driving valve seat magnet 54 (this timewith one-way valve 70 open) back down the length of L-VAD 306 to itsoriginal position. Electromagnetic coils 300, 302 and 304 are energizedby controller 64, in response to the ECG signal 66, and the ejectionvolumes and pressures of the right and left sides of the system asillustrated in FIG. 9, in a similar manner as described in the preferredembodiment.

[0235]FIG. 27 represents an alternative embodiment of the ejectionvolume measuring apparatus for use in an alternative BEEP system,generally designated 35″ (only a portion of which is shown). Inembodiment 35″, instead of rare earth magnets 402, 404, 412, and 414(shown in FIG. 9), any proximity sensor 502, 504, 520 and 522, can besubstituted. Such proximity sensors may be included in any or all of thefour chambers of the heart. These proximity sensors may or may not needto be coupled with conductors 506, 508, 524, and 526. Proximity sensoroutput can be fed to signal converters 512 and 530 directly, or withconductors 510 and 528. The output from signal converter 512 is directedwith lead 408 into wire bundle 410, and the output from signal converter530 is directed with lead 418, also into wire bundle 410. Wire bundle410 can, but does not necessarily, include inside it conductor 63, shownin FIGS. 1-4, which transmits the ECG signal to controller 64. Thesignals from these conductors become inputs to controller 64.

[0236] Optimization Procedure and Control Sequence:

[0237] The following optimization process can be applied to any VAD orTAH device. During normal operation the native heart interacts with theVAD so that the overall system/control responds to X/Xmax, the ECGsignal, and φ of the combined system of the native heart and VAD.

[0238] FIGS. 28-34 illustrate the basic procedure of the optimizationand control process for new BEEP system 35, 35′ or 35″. It is to beunderstood throughout the discussion herein that, unless otherwisespecified, “VAD” shall be interpreted to mean either the L-VAD, R-VAD,BI-VAD or TAH.

[0239] The main steps of the process are the following:

[0240] Develop a mathematical model of the dynamic behavior of thedesired part of the physical system

[0241] Identify the system inputs, outputs and desired constraints forthe physical part of the system in (a) above

[0242] Optimize power input to the VAD to complement the action of thediseased native heart

[0243] Develop an optimized control scheme for the inputs and outputs

[0244] Perform tests on the individual patient

[0245] Maintenance: updating the dynamic optimization and controlschemes as the condition of the patient changes with time.

[0246] The inlet and outlet boundary conditions, and other engineeringinputs to the model, and corresponding dynamic inputs, will depend onthe dynamic system definition (Gyftopoulos and Beretta, 1991). Severalalternative embodiments of the definition of the dynamic system thatwill perform the optimization can be defined, and the following examplesare given to illustrate the flexibility and potential of the poweroptimization method and control optimization method.

[0247]FIG. 28 illustrates the main components of the circulation systemwith the main BI-VAD components in place, with the right and left sideforcing functions from the native heart and the VAD. The level of detailof system components shown in FIG. 28 is for illustration purposes only,and several alternative models with more or fewer details of systemcomponents can be drawn. One could choose to obtain the dynamic responseof the whole system, including theoretical, numerical or experimentalmodels for the pulmonary system (lungs) and oxygenation dynamic systems.A continuous model (differential equations expressing the dynamicresponse at any position and time of the component) for all componentsof the system would be of enormous complexity, though stilltheoretically possible. Continuous models can also lead to distributedparameter analyses. With current (2001) technology those skilled in theart would likely (and straining computational resources) choose tomodel: the L-VAD system from points M1 to M2; the “right side” of theheart system from points M3 to M4 without R-VAD; and from M3 to M5 withthe R-VAD. These M1 to M5 points (or other similar points that may beused for the optimization) are used to separate suitable portions of thephysical system in order to develop a dynamic model of portions of thesystem rather than the whole system shown in FIG. 28. Points such as M1to M5 (or other similar points) used in order to simplify the model, mayrepresent physical cross-sections of the flow passages, usually in themain blood vessels. Those with ordinary skill in the art may choose todevelop the dynamic model using combinations of:

[0248] (a) discretized finite element method programs (FEM, for exampleSzabo and Babuska, 1991; Bathe, 1995). For example, FEM models may beused for the cardiac muscle, for structural mechanical components, andother components;

[0249] (b) computational fluid dynamics models (CFD, for exampleAnderson et al., 1984; Kiris et al., 1997). For example, CFD models maybe used for the blood and hydraulic fluid flows, and other components ofthe system;

[0250] (c) analytic solutions for some dynamic elements. For example,analytic models may be used for some of the fluid leakage in narrowpassages, using lubrication theory, and other boundary-layer techniques.Samples of such methods for different sub-components of the system arepresented by (Nichols and O'Rourke, 1998; Panton, 1984; White, 1991;Schlichting, 1979; Hinze, 1987);

[0251] (d) specialized information for select parts of thecardiovascular system. For example, select such models are described by(Fung, 1984; Braunwald, 1984; Verdonck, 2000; Peskin and McQueen, 1997);

[0252] (e) lumped parameter models for some of the dynamic components.For example, some of the mechanical components may be represented withtechniques described by (Meirovich, 1975); and

[0253] (f) experimental data (which are usually the most reliablemodels) can be used for, any aspect of the dynamic components of thesystem.

[0254] Thus the overall dynamic model can be based on continuummechanics (or their variant of distributed parameter models), can bediscretized, can be based on lumped parameters, rely on experimentaldata, or any combination thereof. Some of the component models will belinear, others will be non-linear, and some will be discrete orpiecewise continuous (for example valve 70 is sometimes open, sometimesclosed, and sometimes in the process of opening or closing). The overalldynamic model is likely to be complex, requiring significantcomputational resources. In alternative embodiments useful informationcan also be derived from simpler piecewise-continuous lumped-parameternon-linear dynamic models for the main components. For example, suchextremely simplified models can consist of several masses, springs anddampers for each of the main components shown in FIG. 28, so that theoverall dynamic model could be run on a conventional desktop personalcomputer.

[0255] Once a suitable part of the physical prototype has been definedwith points such as M1 to M5 for development of the dynamic model, thesuitable inputs, outputs, boundary conditions, and other systemconstraints must be carefully defined (this is referring to the correctsystem definition, mentioned above). FIG. 29 is an example of onepossible dynamic-system representation of the main components of theleft side of the diseased native heart plus the L-VAD. Alternativeembodiments of the model may include portions of the right side of thenative heart and/or the R-VAD. If the sample model for FIG. 29 extendsbetween points M1 and M2 in FIG. 28, then inlet and outlet boundaryconditions would involve combinations of blood pressures and velocitiesat M1 and M2 as functions of time. In FIG. 29 the center diagramillustrates the dynamic model of the physical system. The dynamic modelcan be developed with experimental data or with equations or acombination thereof. There are two basic physical components to thecenter block of FIG. 29. One physical component is the diseased nativeheart and surrounding tissue to the native heart; and the other physicalcomponent is the VAD or BI-VAD or TAH system. The two physicalcomponents interact and dynamically affect each other during normaloperation as shown by the dashed line R. The combination of the two maincomponents of the system, whether presented mathematically or withexperimental data, results in a dynamic representation of the physicalsystem from M1 to M2 that can be described by a form of equation (7).The engineering definitions of the thermodynamic system (dynamic,thermodynamic, fluid dynamic, mechanical, etc., as discussed earlier)are crucial to the analysis and must be such that the patient's tissuesurrounding both the native heart and the VAD or TAH are sufficientlyremoved from the components so that the dynamic operation does notaffect the boundary surface that separates the dynamic system from thesurrounding tissue. This is an imaginary boundary surface typical of theboundary surfaces in thermofluid dynamics texts that define the boundaryof the engineering system being analyzed. The input for the dynamicrepresentation of subsystems and components (shown on the left side ofFIG. 29) can come from different sources. Some of the input can beexperimental data, which is usually the preferred source of data, butother input can come from other adequate mathematical representations.Such inputs may come from the constitutive relationships of cardiacmuscle or other muscle, or of the surrounding tissue, or theconstitutive relationships for blood flow, whether it is modeled asNewtonian or non-Newtonian, incompressible or compressible. Severalillustrations of these models are published in the above references. Themathematical representation of the physical system depicted in FIG. 29from M1 to M2 can use constitutive relationships for the mechanicalcomponents, or constitutive relationships for electromagneticcomponents, or experimental data, or any combination thereof, orexperimental data and mathematical expressions of constitutiverelations, as shown on the left side of FIG. 29.

[0256] Other components of the physical system can also be incorporated,as needed, depending upon how many or how few of the system componentsare used in the dynamic model represented by the final form of “processequations” (7) and (8). The level of detail of the dynamic model willaffect the complexity of the required solution. It will also affect thefidelity and thus accuracy of the results. In general, the higher thecomplexity and the fidelity the more accurate the results, but at somepoint there is a limit; i.e., a point of diminishing returns, where theincreased complexity does not justify the higher accuracy. A judgmentmust be made on the fidelity of the dynamic model required for theparticular application of the invention. The final decision on thisissue will also depend on the sophistication of engineering toolsavailable, (CFD, FEM etc) and the level of accuracy required of theresults.

[0257] There are several alternative types of controllers suitable forthe application. In the simplest case the controllers may give constant(battery) voltage, and vary the currents as a function of time to thethree coils of FIG. 18 (for example i₁{t}, i₂{t} and i₃{t}). Othercontrollers may give constant current but varying voltages. Othercontrollers may vary both the voltage and the current. The latter is themost likely embodiment. The third solution is likely to give the leastpower required than the other two controllers, as the electricalresistance, inductance and impedance of the coils, driving magnet 40,and surrounding ferromagnetic material impose non-linear effects on theunsteady flows of voltage and current, and the third type of controllerallows one to take full advantage of the “natural frequencies” of thedynamic magnetic system in relation to the forcing function required bythe patient-VAD system.

[0258] The dynamic model developed above is used in the poweroptimization method, an example of which is illustrated in FIG. 30. Thepurpose of this optimization method is to minimize power requirementsand maximize battery life between recharges. This is accomplished byidentifying the minimum electrical power required to the coils for eachoperating condition of the VAD. Since the diseased native heart and VADaffect each other's dynamic performance (broken line U in FIG. 29), theoptimization process must be repeated separately for each initialoperating condition of the unaided native heart. The inputs for thespecific illustration example are the initial condition of the diseasednative heart, comprised of the ECG trace and the ejected blood volumesand pressures of the right and left side (atria and ventricles) asfunctions of time over the period of the heart rate. In one embodimentof the optimization process an intermediate output of the poweroptimization method is X/Xmax {t} and F_(vad){t} of the R-VAD and L-VAD.In an alternative embodiment of the optimization process the output isthe voltage and current fed to each coil as a function of time as shownin FIG. 30. This optimization process must be repeated for each initialcondition of the diseased native heart identified above.

[0259] One or more of the several potential optimal solutions V{t} andi{t} are stored in controller 64. The choice of optimal solution toinsert in the controller is illustrated with an example below.

[0260] The example power optimization method of FIG. 30 searches forshapes of X/Xmax{t} that require the minimum power. In one exampleembodiment of this optimization process, the ejection volume of thenative left ventricle is evaluated using MRI, echocardiography, or othersimilar techniques such as correlation with the movement of proximitysensors as described earlier. The desired additional volume that must beprovided by L-VAD 74 is evaluated by methods illustrated in the earliertable of potential sizes of VAD. This dictates the required travelX/Xmax of L-VAD 74 in FIG. 16. Next, an initial estimate for the traceof line 162 in FIG. 17 (starting with an initial shape resembling thatof line 162) is input into the power optimization method. This lineshape may also be modeled with Fourier series analysis, and theamplitudes and phases in these Fourier series have a phase differencefrom the ECG trace of the native heart in F_(nh){t}. One measure ofthese phases is graphically reflected in the phase difference φ from thephase of R in the QRS complex (phase zero) to point 154 in the trace ofX/Xmax in FIG. 17.

[0261] This concept is commonly referred to as “phase” in dynamicsystems. The pressures of the four chambers as functions of time arerequired at least for the optimization sequence, and it may also berequired during the normal running of the device. However, the pressuresand the volumes can also be correlated by other means in normal runningof the device (for examples the ECG signal alone, or the ECG signal plusvolume traces, or ECG signal plus volume plus pressure traces).

[0262] The forcing function of the native heart can be measured (forexample with measurements of ventricular and atrial pressures andvolumes, or their correlations) as described elsewhere in the text. Theforcing function of the VAD is an input to the optimization process asdescribed below. There are several alternative combinations ofspecifying this forcing function as an input to the optimizationprocess. For example, one way is to prescribe the displacement X/Xmax ofdriving magnet 40 as a function of time, (FIGS. 16 and 17), evaluate therequired force on driving magnet 40 from the coils, and then evaluatethe electrical power required from the controller (voltages and/orcurrents to the coils) to accomplish this motion. This is furtherelaborated below. Then in an iterative process the displacement versustime (of driving magnet 40) can be changed until the electrical power tothe coils is minimized while the displacement of driving magnet 40provides corresponding displacements of driven magnet 54 that result inacceptable ranges of volumetric blood throughput and heart rate.

[0263] This initial estimate of the forced motion of X/Xmax {t} resultsin changes in the pressure supplied by the combined diseased leftventricle plus L-VAD 74. The result is that the pressure and volumetraces of the diseased heart shown in FIGS. 14 and 15 are modified,because the dynamic response of the native heart system and L-VAD systemaffect each other. (In simple terms, the motion of the one-way valve 70from the aortic valve to the aortic arch sucks additional blood perheart beat from that accomplished by the diseased native ventriclealone, thus increasing ejected blood volume per beat, so that the wholesystem would tend to operate at lower heart rates).

[0264] The initial estimate of the shape of X/Xmax{t} results in arequired forcing function F_(vad){t} that must be provided by drivingmagnet 40 to the hydraulic fluid and from there to the blood, and thiscorresponds to the distribution of voltage and current over time thatthe coils must provide to the driving magnet 40. The forcing function ondriving magnet 40 is computed using the dynamic model described above inequation (7). This forcing function of the optimized design is comparedwith the 30-36 N maximum force estimated in the discussion of FIGS. 1-4,above. The electrical power (V{t} and i{t}) required to provide thisforce (FIG. 17, line 180).is computed using dynamic models of thetransmission of power from the coils to the magnet, or measuredexperimentally, or with a similar technique, reflecting equation (8).The shape of Xmax {t} versus time is iteratively manipulated until theelectrical power required is minimized.

[0265] In alternative embodiments of the power optimization method thisminimization can be done numerically or experimentally, or by neuralnetworks to handle the volume of data and computations required. Theoptimization of the transmission of electromagnetic power can be donefor at least three different cases, depending on the type of controller64: (a) the coils are supplied with constant voltage and power changesare obtained with changes in the electrical current; (b) The coils aresupplied with constant current and power changes are obtained withchanges in the voltage; and (c) the controller can vary both the currentand voltage applied to each coil as a function of time.

[0266] It is expected that for a given initial diseased heart condition(for example, a heart rate of 100 beats per minute and ejected volumefrom the unaided left ventricle 50 cc) the power optimization processwill result in several combinations of modified heart rates and ejectedvolumes (from the combined left ventricle and L-VAD) with slightlydifferent power requirements. For example, three potential L-VADsolutions, each with a different shape of X/Xmax in FIGS. 16 and 17 tothe above diseased-heart condition, may be:

[0267] a. 80 beats per minute, 80 cc per beat, 7.0 Joules per beat (560Joules/minute);

[0268] b. 130 beats per minute, 60 cc per beat, 3.0 Joules per beat (465Joules/min);

[0269] c. 60 beats per minute, 90 cc per beat, 8.0 Joules per beat (480Joules/min).

[0270] In the last step of FIG. 30, for most practical applications acardiologist would choose to store in controller 64 either solution (a)or solution (c) rather than the lower-energy solution (b). These“optimal” solutions are obtained in an “external optimization process”shown in FIG. 28 for a wide range of diseased heart conditions andstored in controller 64. The combined output of the diseased nativeheart and the new VAD/TAH is optimized both for the individual diseasednative heart of the patient and for the power required to drive theartificial device. The output of the power optimization method is theelectrical power, and combinations of voltage and current, that must beapplied by the controller to the coils.

[0271] In alternative embodiments this power optimization method can becarried out mathematically (equations (7) and (8) for the chosensystem), or experimentally (with patient and VAD) in clinical trials forgroups of patients, or individuals patients. In either case theseoptimization processes would benefit by the use of neural networks.

[0272] The VAD installed in the patient must be able to adapt to dynamicchanges from one condition to the other, as the patient with the VADimplanted in normal operative condition goes through normal dailyactivities requiring changes in heart rates and ejected blood volumes.The optimal design of the multi-input multi-output dynamic system of thepatient is illustrated in FIG. 31. In one alternative embodiment of thecontrol optimization method the physical dynamic system in FIG. 31 isthe patient with the VAD installed, or in other words the controloptimization method is experimental and is done clinically. In anotheralternative embodiment the control optimization method is done with thedynamic model of the physical prototype, reflected in expressionsof“process equations” (7) and (8).

[0273] Examples of the dynamic system shown in FIG. 31 are dynamicmodels such as those shown in FIGS. 28 and 29, and incorporate theresults of the power optimization method of FIG. 30 (that defines thesteady-state, non-dynamically changing conditions of the system). Incontrol-system terminology this is a multivariate control scheme (asopposed to more usual control schemes for simpler linear mechanicalsystems). For purposes of this document, “multivariate” means that theoutput state of the dynamic system is characterized by several inputvariables and several output variables, illustrated by the incoming andoutgoing arrows on the left and right side of FIG. 31.

[0274] Examples of these output-state variables are the ECG trace, theblood volumes ejected from the ventricles, the flow rates through pointssuch as M1, M2, M3 and M5, the blood pressures or hydraulic-fluidpressures at various points in the flow system, other similarquantities, rates of change of these variables with time, orcombinations thereof. The state of these variables is measured byvarious pressure, velocity, position, etc., transducers. Informationabout the output state variables is fed back to the control node (x incircle) via the feedback transfer matrices G(s). The purpose of thecontrol optimization method is:

[0275] (a) to find the optimum output variables for the control scheme;

[0276] (b) to find the types and values of these feedback transfermatrices G(s), which feed back signals to the control node; and

[0277] (c) to find the optimum input state variables for the controlscheme.

[0278] This optimization method is a multivariate input-output controlmethod with several input-output state variables. The input state can bedefined with variables such as beats per minute, the phase φ of thedistance from point Q to point 154 in FIG. 17 (in units of degrees or inunits of time), the value of X/Xmax, the shape of X/Xmax, several othersimilar quantities, or their rates of change with time, and combinationsthereof. In general, these input-state control variables would bedifferent from any similar quantities that were computed in the poweroptimization process. This does not preclude using the quantities fromthe power optimization process, but this may lead to systeminstabilities during transients.

[0279] Thus the inputs to the dynamic system are manipulated byfunctions of the outputs of the dynamic system (which can be thephysical patient and VAD, or the dynamic model developed above) asaffected by the feedback transfer matrices G(s). The modified inputs arefed into the dynamic system and affect its output state. Thisoptimization of the control method can be done analytically,experimentally, or with various heuristic methods. Essentially thesecontrol methods ensure that corrections to the dynamic system toaccommodate transient operations do not become unstable.

[0280] Simplified versions of this control sequence can be analyzed withlinear control theory. However it is more likely that development of thecontrol method will require well established analytic and experimentaltechniques of non-linear, discrete or continuous systems control, asseveral of the elements of the dynamic system (e.g. in FIG. 29) arenon-linear.

[0281] Suitable analytic optimal control techniques have been publishedin the open literature (Brogan, 1990; Glad and Ljung, 2000; Fradkov etal., 1999; Schroder, 2000). However, it is envisioned that inalternative embodiments neural networks, adaptive control techniques,and observer-based methodologies will be suitable alternativeembodiments of the new control optimization method. The final step inthe control optimization sequence is to store the optimized feedbacktransfer matrices G(s), and associated control scheme, into controller64.

[0282] An additional way to illustrate the flow of information flow inthe new device during normal operation is shown in the flow charts ofFIGS. 32, 33 and 34, all of which include points M1, M2, M3 and M5 thatare also in FIGS. 28 and 29. FIG. 32 shows the application to an L-VAD.The figure shows at the top the native heart as right and left sides,atria and ventricles, and in the center there is an illustration of theECG signal, which feeds information to the controller. In this case,shown in FIG. 32, the information would be the ECG trace, or heart rateand phase entering the L-VAD controller. (Alternative embodiments mayinclude measures of volumes and pressures in atria and/or ventricles).The left ventricle provides some output that goes into the controljunction, shown by an X in a circle, which is also fed into thecontroller. The controller, using information provided by the poweroptimization method and control optimization method that were shown anddiscussed with reference to FIGS. 30 and 31, provides electric power,V{t} and i{t} to the coils, that dictates the movement of driving magnet40. Driving magnet 40 in turn dictates the movement of driven magnet 44,and that in turn dictates the movement of valve-seat magnet 54, whichaffects the output of the left ventricle (point M2). A similararrangement could be drawn for the R-VAD, but it is exactly symmetricalto the one shown here so this is not drawn or described further.

[0283] The flow chart of FIG. 33 illustrates the application of the newoptimization process to a bi-ventricular assist device (BI-VAD,described above). The flow chart is split into two parts, for the L-VADand the R-VAD, right and left respectively. The components themselvesand the logic flow paths are similar to those shown in FIG. 32, and thusare not discussed further herein.

[0284]FIG. 34 shows the application of the new optimization process tothe total artificial heart (TAH) embodiment, in which the left and rightventricles are removed, but signals are still received from thesinoatrial node. These signals are provided to the L-TAH and R-TAHcontrollers 64. In response, the controllers drive magnets 40, 44 and54, and finally these magnets provide the overall volumetric throughputfor the cardiac system, corresponding to points M2 and M5 in FIGS. 30and 31, as previously discussed.

[0285] These mathematical and engineering techniques will be augmentedby clinical trials on groups of patients, and standard-sized or uniqueVAD devices sized for individual patients devices may be optimized tothe individual patients. As the condition of the patient changes withtime, the control variables and control scheme stored on controller 64will need to be updated to the new condition of the patient. The powerand control optimization sequences are identical to the sequencesdescribed above. The new data for controller 64 can be transmitted tothe controller inside the patient's body using established infra reddata transmission techniques, or other similar techniques.

[0286] Several alternative embodiments to the described systems andmethods are also conceived. For example, one alternative embodimententails the use of neural networks, or comparable technology, tooptimize the displacements shown in FIG. 28 (X/Xmax and their shape),with the ejected blood volume. For example, with reference to FIG. 29,dynamic measurement of volume ejected and phase may be eliminated,because volume and phase can be correlated (with neural networks) withthe motion of two or more proximity sensors, as shown in FIG. 27.

[0287] Alternative embodiment of the above-described methods areconceived in which the optimization processes for FIGS. 30 and 31 is notdone mathematically, but it largely depends on clinical trials withextensive use of neural networks to expedite the computation process.

[0288] Another enhancement of the new system is that the controller candetect the presence of certain arrhythmias, such as ventriculartachycardia, for example. In this event, the electromagnetic pump coilswould be de-energized and no current would be supplied to such coils, asit would be undesirable for the VAD to be activated. As a “fail-safe”,if the VAD was in fact de-energized, in such a case, the one-way valve70 inside of valve-seat magnet 54 would respond to the pressure gradientof the blood flowing past it, and would open and close via the forcesapplied to it by the flowing blood.

[0289] In view of the foregoing, it will be seen that the severalobjects of the invention are achieved and other advantages are attained.Although the foregoing includes a description of the best modecontemplated for carrying out he invention, various modifications areconceivable. As various modifications could be made in the constructionsand methods herein described and illustrated without departing from thescope of the invention, it is intended that all matter contained in theforegoing description or shown in the accompanying drawings shall beinterpreted as illustrative rather than limiting.

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What is claimed is:
 1. A method of optimizing the power required by a mechanical cardiac pumping device in steady-state operating condition, said method comprising the steps of: a. modeling the physical system, or at least a portion thereof, of the patient who will receive the mechanical cardiac pumping device; b. identifying an operating condition of the native heart of the patient who will receive the mechanical cardiac pumping device to which the mechanical cardiac pumping device will respond; c. using the model of the physical system from step a, above, to determine the required blood volume to be ejected from the mechanical cardiac pumping device; d. providing an initial estimate of the instantaneous power as a function of time across at least one period of the heartbeat required to be provided to the mechanical cardiac pumping device in order to provide the required ejected blood volume; e. evaluating the resultant ejected blood volume with data obtained from the model of the physical system; f. updating the estimate of the power requirement; g. iteratively performing steps e and f, above, until the power required to obtain the required ejected blood volume by the combined operation of the native heart and the VAD is identified; h. determining possible solutions to the instantaneous power as a function of time that allows the mechanical cardiac pumping device to provide the required ejected blood volume; i. choosing the solution from step h, above, that best matches the physiological constraints of the patient and provides for optimal power usage by the mechanical cardiac pumping device; and j. iteratively performing steps b through i, above, until the mechanical cardiac pumping device is optimized to respond to each desired operating condition of the native heart.
 2. A method of optimizing the power and energy required by a mechanical cardiac pumping device in steady-state operating condition, said method comprising the steps of: a. modeling the physical system, or at least a portion thereof, of the patient who will receive the mechanical cardiac pumping device; b. identifying an operating condition of the native heart of the patient who will receive the mechanical cardiac pumping device to which the mechanical cardiac pumping device will respond; c. using the model of the physical system from step a, above, to determine the required blood volume to be ejected from the mechanical cardiac pumping device; d. providing an initial estimate of the instantaneous power as a function of time across at least one period of the heartbeat required to be provided to the mechanical cardiac pumping device in order to provide the required ejected blood volume; e. evaluating the resultant ejected blood volume with data obtained from the model of the physical system; f. updating the estimate of the power requirement; g. iteratively performing steps e and f, above, until the power required to obtain the required ejected blood volume by the combined operation of the native heart and the VAD is identified; h. determining the possible solutions to the instantaneous power as a function of time, and the total energy over the pumping cycle that allows the mechanical cardiac pumping device to provide the required ejected blood volume; i. choosing the solution from step h, above, that best matches the physiological constraints of the patient and provides for optimal power and energy usage by the mechanical cardiac pumping device; and j. iteratively performing steps b through i, above, until the mechanical cardiac pumping device is optimized to respond to each desired operating condition of the native heart.
 3. A method of optimizing the energy required by a mechanical cardiac pumping device in steady-state operating condition, said method comprising the steps of: a. modeling the physical system, or at least a portion thereof, of the patient who will receive the mechanical cardiac pumping device; b. identifying an operating condition of the native heart of the patient who will receive the mechanical cardiac pumping device to which the mechnical cardiac pumping device will respond; c. using the model of the physical system from step a, above, to determine the required blood volume to be ejected from the mechanical cardiac pumping device; d. providing an initial estimate of the instantaneous power as a function of time across at least one period of the heartbeat required to be provided to the mechanical cardiac pumping device in order to provide the required ejected blood volume; e. evaluating the resultant ejected blood volume with data obtained from the model of the physical system; f. updating the estimate of the power requirement; g. iteratively performing steps e and f, above, until the power required to obtain the required ejected blood volume by the combined operation of the native heart and the VAD is identified; h. determining possible solutions to the total energy over the pumping cycle that allows the mechanical cardiac pumping device to provide the required ejected blood volume; i. choosing the solution from step h, above, that best matches the physiological constraints of the patient and provides for optimal energy usage by the mechanical cardiac pumping device; and j. iteratively performing steps b through i, above, until the mechanical cardiac pumping device is optimized to respond to each desired operating condition of the native heart.
 4. A method of optimizing the power required by a mechanical cardiac pumping device in steady-state operating condition, said method comprising the steps of: a. modeling the physical system, or at least a portion thereof, of the patient who will receive the mechanical cardiac pumping device; b. providing an initial estimate of the instantaneous power as a function of time across at least one period of the heartbeat required to be provided to the mechanical cardiac pumping device in order to provide the required ejected blood volume; c. evaluating the resultant ejected blood volume; d. updating the estimate of the power requirement; e. iteratively performing steps c and d, above, until the power required to obtain the required ejected blood volume by the combined operation of the native heart and the VAD is identified; f. determining the possible solutions to the instantaneous power as a function of time that allows the mechanical cardiac pumping device to provide the required ejected blood volume; g. choosing the solution from step f, above, that best matches the physiological constraints of the patient and provides for optimal power usage by the mechanical cardiac pumping device; and h. iteratively performing steps b through g, above, until the mechanical cardiac pumping device is optimized to respond to each desired operating condition of the native heart.
 5. A method of optimizing the energy required by a mechanical cardiac pumping device in steady-state operating condition, said method comprising the steps of: a. modeling the physical system, or at least a portion thereof, of the patient who will receive the mechanical cardiac pumping device; b. providing an initial estimate of the instantaneous power as a function of time across at least one period of the heartbeat required to be provided to the mechanical cardiac pumping device in order to provide the required ejected blood volume; c. evaluating the resultant ejected blood volume; d. updating the estimate of the power requirement; e. iteratively performing steps c and d, above, until the power required to obtain the required ejected blood volume by the combined operation of the native heart and the VAD is identified; f. determining the possible solutions to the total energy over the pumping cycle that allows the mechanical cardiac pumping device to provide the required ejected blood volume; g. choosing the solution from step f, above, that best matches the physiological constraints of the patient and provides for optimal energy usage by the mechanical cardiac pumping device; and h. iteratively performing steps b through g, above, until the mechanical cardiac pumping device is optimized to respond to each desired operating condition of the native heart.
 6. A method of optimizing the power and energy required by a mechanical cardiac pumping device in steady-state operating condition, said method comprising the steps of: a. modeling the physical system, or at least a portion thereof, of the patient who will receive the mechanical cardiac pumping device; b. providing an initial estimate of the instantaneous power as a function of time across at least one period of the heartbeat required to be provided to the mechanical cardiac pumping device in order to provide the required ejected blood volume; c. evaluating the resultant ejected blood volume; d. updating the estimate of the power requirement; e. iteratively performing steps c and d, above, until the power required to obtain the required ejected blood volume by the combined operation of the native heart and the VAD is identified; f. determining the possible solutions to the instantaneous power as a function of time and total energy over the pumping cycle that allows the mechanical cardiac pumping device to provide the required ejected blood volume; g. choosing the solution from step f, above, that best matches the physiological constraints of the patient and provides for optimal power and energy usage by the mechanical cardiac pumping device; and h. iteratively performing steps b through g, above, until the mechanical cardiac pumping device is optimized to respond to each desired operating condition of the native heart.
 7. The method of any one of claims 1 through 6, wherein the modeled physical system is used to determine the required blood volume to be ejected from the mechanical cardiac pumping device by discretizing the modeled physical system with Finite Element Methods and Computational Fluid Dynamics into mass, damping, and stiffness matrices, and their corresponding elemental displacements.
 8. The method of any one of claims 1 through 6, wherein the physical system is modeled with MRI data and the modeled physical system is used to determine the required ejected blood volume from the mechanical cardiac pumping device by evaluating the MRI data.
 9. The method of any one of claims 1 through 6, wherein at least some components of the physical system are modeled utilizing a lumped-parameter model.
 10. The method of any one of claims 1 through 6, wherein at least some components of the physical system are modeled utilizing a distributed-parameter model.
 11. The method of any one of claims 1 through 6, wherein at least some components of the physical system are modeled utilizing a continuum model.
 12. The method of any one of claims 1 through 6, wherein neural networks are utilized to determine the instantaneous power as a function of time, and the total energy over the pumping cycle, that allows the cardiac device to provide the required ejected blood volume.
 13. The method of any one of claims 1 through 6, wherein heuristic methods are utilized to determine the instantaneous power as a function of time, and the total energy over the pumping cycle, that allows the cardiac device to provide the required ejected blood volume.
 14. The method of any one of claims 1 through 6, wherein the operating condition of the native heart to which the mechanical cardiac pumping device will respond is heart rate.
 15. The method of any one of claims 1 through 6, wherein the operating condition of the native heart to which the mechanical cardiac pumping device will respond is ventricular volume.
 16. The method of any one of claims 1 through 6, wherein the operating condition of the native heart to which the mechanical cardiac pumping device will respond is ventricular pressure.
 17. The method of any one of claims 1 through 6, wherein the operating condition of the native heart to which the mechanical cardiac pumping device will respond is at least a portion of the ECG signal.
 18. The method of any one of claims 1 through 6, wherein the mechanical cardiac pumping device is a left ventricular-assist device and the operating conditions of the native heart to which the device will respond are at least one of: heart rate, heart phase, left ventricular volume, right ventricular volume, left ventricular pressure, and right ventricular pressure.
 19. The method of any one of claims 1 through 6, wherein the mechanical cardiac pumping device is a left ventricular-assist device and the operating conditions of the native heart to which the device will respond are at least one of: heart rate as a function of time, heart phase, left ventricular volume as a function of time, right ventricular volume as a function of time, left ventricular pressure as a function of time, and right ventricular pressure as a function of time.
 20. The method of any one of claims 1 through 6, wherein the mechanical cardiac pumping device is a right ventricular-assist device and the operating conditions of the native heart to which the device will respond are at least one of: heart rate, heart phase, left ventricular volume, right ventricular volume, left ventricular pressure, and right ventricular pressure.
 21. The method of any one of claims 1 through 6, wherein the mechanical cardiac pumping device is a right ventricular-assist device and the operating conditions of the native heart to which the device will respond are at least one of: heart rate as a function of time, heart phase, left ventricular volume as a function of time, right ventricular volume as a function of time, left ventricular pressure as a function of time, and right ventricular pressure as a function of time.
 22. The method of any one of claims 1 through 6, wherein the mechanical cardiac pumping device is a bi-ventricular-assist device and the operating conditions of the native heart to which the device will respond are at least one of: heart rate, heart phase, left ventricular volume, right ventricular volume, left ventricular pressure, and right ventricular pressure.
 23. The method any one of claims 1 through 6, wherein the mechanical cardiac pumping device is a bi-ventricular-assist device and the operating conditions of the native heart to which the device will respond are at least one of: heart rate as a function of time, heart phase, left ventricular volume as a function of time, right ventricular volume as a function of time, left ventricular pressure as a function of time, and right ventricular pressure as a function of time.
 24. The method of any one of claims 1 through 6, wherein the mechanical cardiac pumping device is a total artificial heart.
 25. A method of optimizing the control scheme of a controller for a mechanical cardiac pumping device, said method comprising the steps of: a. simulating the steady-state physical condition of a patient who will receive the mechanical cardiac pumping device; b. identifying a new, target steady-state condition; c. determining which outputs of the physical system to monitor to best perform the transition from one steady-state to another; d. determining the best combination of inputs, outputs, and modifications that achieve transient transfer from one steady-state to another without destabilizing the dynamic system; and e. storing the information determined in steps c and d, above, in the controller.
 26. A method of assisting the cardiac function of the native heart of a patient using an implanted ventricular-assist device, said method comprising the steps of: a. monitoring with a controller a steady-state condition of the physical system of the patient having the implanted ventricular-assist device; b. directing with the controller the optimal power required to sustain the steady-state while meeting physiological constraints; c. determining with the controller when the physical system of the patient has left the steady-state operating condition; d. determining with a controller a new, target steady-state condition; e. while the physical system of the is not in a steady-state operating condition, iteratively performing the following steps i-v: i. monitoring inputs from the physical system of the patient, the inputs being at least one of: a measure of the heart phase, X/Xmax, and the shape of X/Xmax; ii. evaluating with the controller the desired outputs from the combined native heart and ventricular-assist device required at a new steady-state condition, the outputs being at least one of: heart rate, blood volume ejected by the native heart, blood volume ejected by the ventricular-assist device, and the ECG trace; iii. monitoring with a controller the actual outputs of the physical dynamic system, the outputs being at least one of: heart rate, blood volume ejected by the native heart, blood volume ejected by the ventricular-assist device, and at least a portion of the ECG trace; iv. modifying with a controller the actual output data according to feedback transfer matrices stored within the controller; v. transmitting with the controller modified inputs from step iv, above, such that the desired outputs from step ii, above, are achieved without destabilizing the dynamic system of the patient during the transient period between steady-states; f. iteratively performing the steps a-e, above, so long as the ventricular-assist device is in operation.
 27. A method of assisting the cardiac function of the native heart of a patient using an implanted total artificial heart device, said method comprising the steps of: a. monitoring with a controller a steady-state condition of the physical system of the patient having the implanted total artificial heart device; b. directing with the controller the minimum power required to sustain the steady-state while meeting physiological constraints; c. determining with the controller when the physical system of the patient has left the steady-state operating condition; d. determining with a controller a new, target steady-state condition; e. while the physical system of the is not in a steady-state operating condition, iteratively performing the following steps i-v: i. monitoring inputs from the physical system of the patient, the inputs being at least one of: a measure of the heart phase, X/Xmax, and the shape of X/Xmax; ii. evaluating with the controller the desired outputs from the combined native heart and ventricular-assist device required at a new steady-state condition, the outputs being at least one of: blood volume ejected by the total artificial heart, and at least a portion of the ECG trace; iii. monitoring with a controller the actual outputs of the physical dynamic system, the outputs being at least one of: blood volume ejected by the total artificial heart, and at least a portion of the ECG trace; iv. modifying with a controller the actual output data according to feedback transfer matrices stored within the controller; v. transmitting with the controller modified inputs from step iv, above, such that the desired outputs from step ii, above, are achieved without destabilizing the dynamic system of the patient during the transient period between steady-states; f. iteratively performing the steps a-e, above, so long as the total artificial heart device is in operation.
 28. A method of assisting the cardiac function of the native heart of a patient using an implanted ventricular-assist device, said method comprising the steps of: a. allowing the native heart to pump as much blood as it is able prior to activation of the ventricular-assist device; b. activating the ventricular-assist device to provide additional pumping action as the blood-ejection phase of the native heart nears completion such that the native heart pumps more blood than it would unaided due to the reduction of back pressure in the native ventricle caused by the pumping action of the ventricular-assist device; c. coordinating the timing of the action and length of the pumping stroke of the ventricular-assist device with the ejected blood volume and rhythm of the native heart such that the power required by the ventricular-assist device is the optimal needed to pump the required volume of blood while meeting physiological constraints; d. varying the stroke displacement over time and resulting power over time of the ventricular-assist device such that the power required by the ventricular-assist device is the optimal needed to pump the required volume of blood while meeting physiological constraints; e. iteratively performing steps a through d, above, so long as the ventricular-assist device is in operation.
 29. A method of optimizing a mechanical cardiac pumping device wherein unsteady fluid mechanics are used to optimize the forcing function imposed by the mechanical cardiac pumping device such that the power required by the mechanical cardiac pumping device is the minimum power required to complement the cardiac output of the diseased native heart, said method comprising the steps of: a. modeling the dynamic response of the diseased native heart and of the mechanical cardiac pumping device with experimental data; b. using the instantaneous non-linear mass, [M], damping, [C], and stiffness, [K] matrices of the dynamic model, and corresponding elemental displacements {x} and its derivatives {x.} and {x}, as inputs into an equation which sums these matrices to calculate the forcing function, F{t}, of the dynamic system; c. calculating the forcing function of the diseased native heart, F_(nh){t}; d. calculating the required forcing function of the mechanical cardiac pumping device, F_(vad){t}; e. inputing the value of F_(vad){t}from step d, above, into a controller; and f. connecting operatively the controller to a mechanical cardiac pumping device, such that the controller is able to direct to the mechanical cardiac pumping device the minimum power required to achieve F_(vad){t}.
 30. A method of optimizing a mechanical cardiac pumping device wherein unsteady fluid mechanics are used to optimize the forcing function imposed by the mechanical cardiac pumping device such that the power required by the mechanical cardiac pumping device is the minimum power required to complement the cardiac output of the diseased native heart, said method comprising the steps of: a. modeling the dynamic response of the diseased native heart and of the mechanical cardiac pumping device with experimental data; b. using the instantaneous non-linear mass, [M], damping, [C], and stiffness, [K] matrices of the dynamic model, and corresponding elemental displacements {x} and its derivatives {{umlaut over (x)}} and {{dot over (x)}}, as inputs into an equation of the form: [M]{{umlaut over (x)}}+[C]{{dot over (x)}}+[K]{x}=F{t} to calculate the forcing function, F{t}, of the dynamic system; c. calculating the forcing function of the diseased native heart, F_(nh){t}; d. calculating the required forcing function of the mechanical cardiac pumping device, F_(vad){t}, using an equation of the form: F{t}=F _(nh) {t}+F _(vad) {t} e. inputing the value of F_(vad){t} from step d, above, into a controller; and f. connecting operatively the controller to a mechanical cardiac pumping device, such that the controller is able to direct to the mechanical cardiac pumping device the optimal power required to achieve F_(vad){t}.
 31. A method of optimizing a mechanical cardiac pumping device wherein unsteady fluid mechanics are used to optimize the forcing function imposed by the mechanical cardiac pumping device such that the power required by the mechanical cardiac pumping device is the minimum power required to complement the cardiac output of the diseased native heart, said method comprising the steps of: a. modeling the dynamic response of the diseased native heart and of the mechanical cardiac pumping device with experimental data; b. using the instantaneous non-linear mass, [M], damping, [C], and stiffness, [K] matrices of the dynamic model, and corresponding elemental displacements {x} and its derivatives {{umlaut over (x)}} and {{dot over (x)}}, as inputs into an equation of the form: [M]{{umlaut over (x)}}+[C]{{dot over (x)}}+[K]{x}=F{t} to calculate the forcing function, F{t}, of the dynamic system; c. calculating the forcing function of the diseased native heart, F_(nh){t}; d. calculating the required forcing function of the mechanical cardiac pumping device, F_(vad){t}, using an equation of the form: F{t}=F _(nh) {t}+F _(vad) {t} e. balancing the instantaneous power at any time t utilized by the mechanical cardiac pumping device with an equation of the form: W(t)=F _(vad) {t}·{{dot over (x)}}+losses=V{t}·i{t} f. inputing the value W(t) from step e, above, into a controller; and g. connecting operatively the controller to a mechanical cardiac pumping device, such that the controller is able to direct to the mechanical cardiac pumping device the optimal power required to achieve F_(vad){t}.
 32. A device to assist the function of a cardiac ventricle, the device comprising: a. a first magnet having an open center and formed of high ferromagnetic-constant material; b. a first vessel surrounding the first magnet and defining a space in fluid communication with the blood flow output great vessel of the diseased ventricle of the patient using the device, the first magnet being movable within the first vessel in substantially fluid-tight relation thereto; c. a second magnet formed of high ferromagnetic-constant material and in magnetic communication with the first magnet so that the respective magnetic fluxes of the first magnet and the second magnet affect each other, so that the first magnet and the second magnet are biased toward and tend to lock to one another, to thereby move in the same direction as one another; d. a second vessel encasing the second magnet and defining a space, the second magnet being movable within the space in substantially fluid-tight relation to the second vessel, the space defined by the second vessel being in fluid communication with a hydraulic pump for actuating the second magnet; and e. an one-way valve connected to the first magnet, the one-way valve being movable with the first magnet, and adapted to cause movement of blood from the diseased ventricle to and into the great vessel associated with that diseased ventricle.
 33. The device of claim 32, wherein the device is a L-VAD and is sized and shaped for positioning between the aortic valve and the aortic arch of the patient using the device.
 34. The device of claim 32, wherein the device is a R-VAD and is sized and shaped for positioning between the pulmonary valve and the bifurcation of the pulmonary trunk of the patient using the device.
 35. A device (58, 74) to assist the function of a cardiac ventricle, the device comprising: a. a first annular magnet (54) formed of high ferromagnetic-constant material; b. a first sleeve (68) surrounding the first annular magnet (54) and defining a space in fluid communication with the blood flow output great vessel of the patient using the device, the first annular magnet (54) being longitudinally and reciprocally slideable within the first sleeve in substantially fluid-tight relation thereto; c. a second annular magnet (44) formed of high ferromagnetic-constant material and sized and shaped for placement exterior of the first sleeve (68), the second annular magnet (44) being disposed coaxially in relation to and in magnetic communication with the first annular magnet (54), so that the respective magnetic fluxes of the first magnet and the second magnet affect each other, so that the first annular magnet and the second annular magnet are biased toward and tend to lock to one another, and to thereby move in the same direction as one another; d. a second sleeve (72) encasing the second annular magnet (44), the second annular magnet being longitudinally and reciprocally slideable between the first sleeve (68) and the second sleeve in substantially fluid-tight relation to the first sleeve and the second sleeve, and the second sleeve (72) defining an annular space (86) radially outwardly of the first sleeve (68) for longitudinal travel therein of the second annular magnet (44), the annular space (86) being in fluid communication with a hydraulic pump for actuating the second annular magnet (44); and e. an one-way valve (70) connected to the first annular magnet (54) and disposed transversely in relation to the longitudinal axis of the first annular magnet (54), the one-way valve being movable with the first magnet, and closed when moving away from the origin of the valve annulus when the device is in normal use position, to thereby cause blood of the patient to move out of a diseased ventricle and toward the great vessels associated with that diseased ventricle as the first annular magnet (54) moves in a direction toward the great vessels of the diseased ventricle due to magnetic flux of the second annular magnet, the one-way valve further being adapted to be open when moving in a direction away from the great vessels of the diseased ventricle, to thereby permit blood of the patient to flow through the one-way valve into the space defined by the first sleeve (68) when the second annular magnet moves away from the great vessels of the diseased ventricle.
 36. The device of claim 35, wherein the entire device (58, 74) is of sufficiently small size and weight for placement in normal operative position between the valve annulus and at least a portion of the great vessels of the diseased ventricle of a patient using the device, to assist or replace the function of at least a portion of the diseased native heart.
 37. The device of claim 35, wherein the device is a L-VAD and is sized and shaped for positioning between the aortic valve and the aortic arch of the patient using the device.
 38. The device of claim 35, wherein the device is a R-VAD and is sized and shaped for positioning between the pulmonary valve and the bifurcation of the pulmonary trunk of the patient using the device.
 39. A system for assisting cardiac ventricular function, the system comprising a hydraulic pumping assembly and a cardiac ventricular assist device in fluid communication with the hydraulic pumping assembly, wherein the ventricular assist device comprises: a. a first magnet having an open center and formed of high ferromagnetic-constant material; b. a first vessel surrounding the first magnet and defining a space in fluid communication with the blood flow output great vessel of the diseased ventricle of the patient using the device, the first magnet being movable within the first vessel in substantially fluid-tight relation thereto; c. a second magnet formed of high ferromagnetic-constant material and being in magnetic communication with the first magnet, so that the respective magnetic fluxes of the first magnet and the second magnet affect each other, so that the first magnet and the second magnet are biased toward and tend to lock to one another, to thereby move in the same direction as one another; d. a second vessel encasing the second magnet and defining a space, the second magnet being movable within the space in substantially fluid-tight relation to the second vessel, the space defined by the second vessel being in fluid communication with a hydraulic pump for actuating the second magnet; and e. an one-way valve connected to the first magnet, the one-way valve being movable with the first magnet, and adapted to cause movement of blood from the diseased ventricle to and into the great vessel associated with that diseased ventricle.
 40. The system of claim 39, wherein the ventricular assist device is a L-VAD and is sized and shaped for positioning between the aortic valve and the aortic arch of the patient using the device.
 41. The device of claim 39, wherein the ventricular assist device is a R-VAD and is sized and shaped for positioning between the pulmonary valve and the bifurcation of the pulmonary trunk of the patient using the device.
 42. A system for assisting cardiac ventricular function, the system comprising a hydraulic pumping assembly and a cardiac ventricular assist device in fluid communication with the cardiac ventricular assist device, wherein the ventricular assist device comprises: a. a first annular magnet (54) formed of high ferromagnetic-constant material; b. a first sleeve (68) surrounding the first annular magnet (54) and defining a space in fluid communication with the blood flow output great vessel of the diseased ventricle of the patient using the device, the first annular magnet (54) being longitudinally and reciprocally slideable within the first sleeve in substantially fluid-tight relation thereto; c. a second annular magnet (44) formed of high ferromagnetic-constant material and sized and shaped for placement exterior of the first sleeve (68), the second annular magnet (44) being disposed coaxially in relation to and in magnetic communication with the first annular magnet (54), so that the respective magnetic fluxes of the first magnet and the second magnet affect each other, so that the first annular magnet and the second annular magnet are biased toward and tend to lock to one another, and to thereby move in the same direction as one another; d. a second sleeve (72) encasing the second annular magnet (44), the second annular magnet being longitudinally and reciprocally slideable between the first sleeve (68) and the second sleeve in substantially fluid-tight relation to the first sleeve and the second sleeve, and the second sleeve (72) defining an annular space (86) radially outwardly of the first sleeve (68) for longitudinal travel therein of the second annular magnet (44), the annular space (86) being in fluid communication with a hydraulic pump for actuating the second annular magnet (44); and e. an one-way valve (70) connected to the first annular magnet (54) and disposed transversely in relation to the longitudinal axis of the first annular magnet (54), the one-way valve being movable with the first magnet, and closed when moving away from the origin of the valve annulus when the device is in normal use position, to thereby cause blood of the patient to move out of a diseased ventricle and toward the great vessels associated with that diseased ventricle as the first annular magnet (54) moves in a direction toward the great vessels of the diseased ventricle due to magnetic flux of the second annular magnet, the one-way valve further being adapted to be open when moving in a direction away from the great vessels of the diseased ventricle, to thereby permit blood of the patient to flow through the one-way valve into the space defined by the first sleeve (68) when the second annular magnet moves away from the great vessels of the diseased ventricle.
 43. The system of claim 42, wherein the entire device (58, 74) is of sufficiently small size and weight for placement in normal operative position between the valve annulus and the great vessels of the diseased ventricle of a patient using the device, to assist or replace the function of at least a portion of the diseased native heart.
 44. A system for assisting cardiac ventricular function, the system comprising a hydraulic pumping assembly and a cardiac ventricular assist device (VAD) in fluid communication with the hydraulic pumping assembly, wherein the hydraulic pumping assembly comprises: a. an encapsulated hydraulic pump having: a pumping chamber for retaining hydraulic fluid therein, the pumping chamber having opposed first and second ends; at least one electromagnetic coil surrounding the pumping chamber; a substantially solid high ferromagnetic-constant magnet disposed longitudinally, slideably and reciprocally within the pumping chamber to act as a piston for driving hydraulic fluid within the pumping chamber in response to signals from a battery/controller assembly; b. a fluid line having a first end and a second end, the first end of the fluid line being connected to and in fluid communication with an the first end of the pumping chamber and the second end of the fluid line being connected to and in fluid communication with the second end of the pumping chamber, the VAD being in fluid communication with the fluid line at a point on the fluid line after the point of connection of the check valve and before the connection of the second end of the fluid line and the second end pump chamber; and c. a battery/controller assembly operatively connected to the check valve and to the at least one electromagnetic coil to provide electric power and control signals to the pump, the battery controller assembly in electrical communication with the native heart of the patient using the system, to thereby receive signals corresponding to physiological parameters from the native heart.
 45. The system of claim 44, wherein the hydraulic pumping assembly further comprises a first end cap and a second end cap connected at opposed first and second ends of the pumping chamber, the first end cap and the second end cap each having an aperture in fluid communication with a hydraulic fluid line.
 46. The system of claim 45, and further comprising a check valve operatively connected in the fluid line between the connection of the first end of the fluid line and the second end of the fluid line to the first and second ends of the pumping chamber, respectively and the fluid line being in fluid communication with the VAD, after the point of connection of the check valve and before the connection of the second end of the fluid valve and the second end cap of the pump cylinder.
 47. The system of claim 45, wherein the information received from the native heart by the battery/controller assembly is at least a portion of an ECG signal from the patient.
 48. The system of claim 45, wherein the information received from the,native heart by the battery/controller assembly is blood pressure information.
 49. The system of claim 45, wherein the information received from the native heart by the battery/controller assembly is blood volume information.
 50. The system of claim 45, wherein the at least one electromagnet coil is three electromagnetic coils disposed longitudinally and coaxially adjacent to one another along the length of the hydraulic pump.
 51. A system for assisting cardiac ventricular function, the system comprising a hydraulic pumping assembly and a cardiac ventricular assist device (VAD) in fluid communication with the hydraulic pumping assembly, wherein the hydraulic pumping assembly comprises: a. a hydraulic pump (42) having: at least one electromagnetic coil (46, 48, 50) encapsulated so as to be fluid-tight, and defining a pumping chamber for retaining hydraulic fluid (52) therein, the pumping chamber having first and second opposed ends; a first end cap (56) and a second end cap (57) connected at opposed first and second ends of the pumping chamber, respectively, the first end cap and the second end cap each having an aperture in fluid communication with a hydraulic fluid line, a substantially solid high ferromagnetic-constant magnet (40) disposed longitudinally, slideably and reciprocally within the pumping chamber to act as a piston for driving hydraulic fluid within the pumping chamber in response to signals from a batter/controller assembly; b. a fluid line (59, 60)having a first end and a second end, the first end of the fluid line being connected to and in fluid communication with an aperture in the first end cap and the second end of the fluid line being connected to and in fluid communication with an aperture in the second end cap, and the VAD being in fluid communication with the hydraulic pumping assembly at a point on the fluid line between the first end and the second end of the fluid line; c. a check valve (84) operatively connected in the fluid line between the connection of the first end of the fluid line and the second end of the fluid line to the first and second end caps respectively, and the fluid line being in fluid communication with the VAD, after the point of connection of the check valve and before the connection of the second end of the fluid valve and the second end cap of the pump cylinder; d. a battery/controller assembly (65) operatively connected to the check valve and to the at least one electromagnetic coil to provide electric power and control signals to the pump, the battery controller assembly in electrical communication with the native heart of the patient using the system, to thereby receive electronic information, including at least portions of ECG signals, blood pressure signals and/or blood volume signals, from the native heart.
 52. The system of claim 51, wherein the battery controller assembly and the hydraulic pump are of sufficiently small size and weight to be entirely contained within the abdominal cavity of the patient using the system and the VAD is of sufficiently small size and weight to be entirely contained within the chest cavity of the patient using the system, and the complete system, including all wires and hydraulic fluid lines, is entirely contained within the body of the patient using the system, so that there is no part of the system extending exterior of the skin of a patient using the system when the system is in normal use position in the patient.
 53. A system for assisting cardiac ventricular function, the system comprising: a hydraulic pumping assembly and a cardiac ventricular assist device in fluid communication with the cardiac ventricular assist device, wherein the ventricular assist device comprises: a. a first annular magnet (54) formed of high ferromagnetic-constant material; b. a first sleeve (68) surrounding the first annular magnet (54) and defining a space in fluid communication with the blood flow output great vessel of a diseased ventricle of the patient using the device, the first annular magnet (54) being longitudinally and reciprocally slideable within the first sleeve in substantially fluid-tight relation thereto; c. a second annular magnet (44) formed of high ferromagnetic-constant material and sized and shaped for placement exterior of the first sleeve (68), the second annular magnet (44) being disposed coaxially in relation to and in magnetic communication with the first annular magnet (54), so that the respective magnetic fluxes of the fist magnet and the second magnet affect each other, so that the first annular magnet and the second annular magnet are biased toward and tend to lock to one another, and to thereby move in the same direction as one another; d. a second sleeve (72) encasing the second annular magnet (44), the second annular magnet being longitudinally and reciprocally slideable between the first sleeve (68) and the second sleeve in substantially fluid-tight relation to the first sleeve and the second sleeve, and the second sleeve (72) defining an annular space (86) radially outwardly of the first sleeve (68) for longitudinal travel therein of the second annular magnet (44), the annular space (86) being in fluid communication with a hydraulic pump for actuating the second annular magnet (44); and e. an one-way valve (70) connected to the first annular magnet (54) and disposed transversely in relation to the longitudinal axis of the first annular magnet (54), the one-way valve being movable with the first magnet, and closed when moving away from the origin of the valve annulus when the device is in normal use position, to thereby cause blood of the patient to move out of a diseased ventricle and toward the great vessels associated with that diseased ventricle as the first annular magnet (54) moves in a direction toward the great vessels of the diseased ventricle due to magnetic flux of the second annular magnet, the one-way valve further being adapted to be open when moving in a direction away from the great vessels of the diseased ventricle, to thereby permit blood of the patient to flow through the one-way valve into the space defined by the first sleeve (68) when the second annular magnet moves away from the great vessels of the diseased ventricle; and further wherein the hydraulic pumping assembly comprises: a. a hydraulic pump (42) having: at least one electromagnetic coil (46, 48, 50) encapsulated so as to be fluid-tight, and defining a pumping chamber for retaining hydraulic fluid (52) therein, the pumping chamber having first and second opposed ends; a first end cap (56) and a second end cap (57) connected at opposed first and second ends of the pumping chamber, respectively, the first end cap and the second end cap each having an aperture in fluid communication with a hydraulic fluid line, a substantially solid high ferromagnetic-constant magnet (40) disposed longitudinally, slideably and reciprocally within the pumping chamber to act as a piston for driving hydraulic fluid within the pumping chamber in response to signals from a batter/controller assembly; b. a fluid line (59, 60) having.a first end and a second end, the first end of the fluid line being connected to and in fluid communication with an aperture in the first end cap and the second end of the fluid line being connected to and in fluid communication with an aperture in the second end cap, and the VAD being in fluid communication with the hydraulic pumping assembly at a point on the fluid line between the first end and the second end of the fluid line; c. a check valve (84) operatively connected in the fluid line between the connection of the first end of the fluid line and the second end of the fluid line to the first and second end caps respectively, and the fluid line being in fluid communication with the VAD, after the point of connection of the check valve and before the connection of the second end of the fluid valve and the second end cap of the pump cylinder; and d. a battery/controller assembly (65) operatively connected to the check valve and to the at least one electromagnetic coil to provide electric power and control signals to the pump, the battery controller assembly in electrical communication with the native heart of the patient using the system, to thereby receive electronic information, including at least portions of ECG signals, blood pressure signals and/or blood volume signals, from the native heart.
 54. The system of claim 53, wherein the entire device (58, 74) is of sufficiently small size and weight for placement in normal operative position between the valve annulus and the great vessels of the diseased ventricle of a patient using the device, to assist or replace the function of at least a portion of the diseased native heart.
 55. The system of claim 53, wherein the battery controller assembly and the hydraulic pump are of sufficiently small size and weight to be entirely contained within the abdominal cavity of the patient using the system and the VAD is of sufficiently small size and weight to be entirely contained within the chest cavity of the patient using the system, and the complete system, including all wires and hydraulic fluid lines, is entirely contained within the body of the patient, so that there is no part of the system extending exterior of the skin of a patient using the system when the system is in normal use position in the patient.
 56. A BI-VAD assembly to assist the function of both the right and left cardiac ventricles simultaneously, the BI-VAD assembly comprising: a. a L-VAD disposed between the aortic valve and the aortic arch, to thereby permit blood to move from the left ventricle of the native heart through the aortic valve and into the L-VAD of a patient using the system, the L-VAD pumping blood into the aortic arch; and b. a R-VAD disposed between the pulmonary valve and the bifurcation of the pulmonary trunk in normal use position in a patient using the BI-VAD, to thereby permit blood to move from the right ventricle of the patient through the pulmonary valve and into the R-VAD as the R-VAD pumps blood into the bifurcation of the pulmonary arteries of the patient.
 57. A BI-VAD assembly (77) to assist the function of both the right and left cardiac ventricles simultaneously, the BI-VAD assembly comprising: a. a L-VAD and a R-VAD; b. the L-VAD being sized and shaped for positioning between the aortic valve and the aortic arch of the patient using the device; the L-VAD comprising: a. a first magnet having an open center and formed of high ferromagnetic-constant material; b. a first vessel surrounding the first magnet and defining a space in fluid communication with the aortic arch of the patient using the device, the first magnet being movable within the first vessel in substantially fluid-tight relation thereto; c. a second magnet formed of high ferromagnetic-constant material and being in magnetic communication with the first magnet, so that the respective magnetic fluxes of the first magnet and the second magnet affect each other, so that the first magnet and the second magnet are biased toward and tend to lock to one another, to thereby move in the same direction as one another; d. a second vessel encasing the second magnet and defining a space, the second magnet being movable within the space in substantially fluid-tight relation to the second vessel, the space defined by the second vessel being in fluid communication with a hydraulic pump for actuating the second magnet; and e. an one-way valve connected to the corresponding first magnet of the L-VAD, the one-way valve being movable with the first magnet of the L-VAD and closed when moving in a direction toward the aortic arch when the device is in normal use position in a patient using the device, to thereby cause blood of the patient to push through the aortic arch as the first magnet moves toward the aortic arch of the patient when the second magnet is actuated to move toward the aortic arch, the one-way valve in the L-VAD further being open when moving away from the aortic arch, to thereby permit blood of the patient to flow through the one-way valve of the L-VAD into the space defined by the first vessel when the second magnet of the L-VAD is moved away from the aortic arch; and c. the R-VAD being sized and shaped for positioning between the pulmonary valve and the bifurcation of the pulmonary trunk of the patient using the device and connected to the L-VAD; the R-VAD comprising: a. a first magnet having an open center and formed of high ferromagnetic-constant material; b. a first vessel surrounding the first magnet and defining a space in fluid communication with the bifurcation of the pulmonary arteries of the patient using the device, the first magnet being movable within the first vessel in substantially fluid-tight relation thereto; c. a second magnet formed of high ferromagnetic-constant material and being in magnetic communication with the first magnet, so that the first magnet and the second magnet are biased toward and tend to lock to one another, to thereby move in the same direction as one another; d. a second vessel encasing the second magnet and defining a space, the second magnet being movable within the space in substantially fluid-tight relation to the second vessel, the space defined by the second vessel being in fluid communication with a hydraulic pump for actuating the second magnet; and e. an one-way valve being movable with the first magnet of the R-VAD and closed when moving toward the bifurcation of the pulmonary arteries when the R-VAD is in normal use position in a patient using the assembly, to thereby cause blood of the patient to push through the bifurcation of the pulmonary arteries as the first magnet of the R-VAD moves toward such bifurcation when the second magnet of the R-VAD is actuated to move toward the bifurcation, the one-way valve of the R-VAD further being open when moving away from the bifurcation of the pulmonary arteries, to thereby permit blood of the patient to flow through the one-way valve of the R-VAD into the space defined by the first vessel when the second magnet of the R-VAD is moved away from the bifurcation of the pulmonary arteries.
 58. A system for assisting cardiac ventricular function simultaneously in both diseased ventricles of the native heart of a.patient using the system, the system comprising at least one hydraulic pumping assembly and two cardiac ventricular assist devices in fluid communication with the at least one hydraulic pumping assembly, wherein the ventricular assist devices comprise: a. a L-VAD disposed between the aortic valve and the aortic arch of the patient, to thereby permit blood to move from the left ventricle of the native heart through the aortic valve and into the L-VAD in a patient using the system, the L-VAD pumping blood into the aortic arch; and b. a R-VAD disposed between the pulmonary valve and bifurcation of the pulmonary trunk in normal use position in a patient using the BI-VAD, to thereby permit blood to move from the right ventricle of the patient through the pulmonary valve and into the R-VAD as the R-VAD pumps blood into the bifurcation of the pulmonary arteries of the patient; and wherein the at least one hydraulic pumping assembly comprises: a. a hydraulic pump having: an encapsulated pumping chamber for retaining hydraulic fluid therein, the pumping chamber having opposed first and second ends; at least one electromagnetic coil surrounding the pumping chamber; and a substantially solid high ferromagnetic-constant magnet disposed longitudinally, slideably and reciprocally within the pumping chamber to act as a piston for driving hydraulic fluid within the pumping chamber in response to signals from a battery/controller assembly; b. a fluid line having a first end and a second end, the first end of the fluid line being connected to and in fluid communication with an the first end of the pumping chamber and the second end of the fluid line being connected to and in fluid communication with the second end of the pumping chamber, the L-VAD and the R-VAD being in fluid communication with the fluid line at a point on the fluid line after the point of connection of the check valve and before the connection of the second end of the fluid line and the second end pump chamber; and c. a battery/controller assembly operatively connected to the check valve and to the at least one electromagnetic coil to provide electric power and control signals to the pump, the battery controller assembly in electrical communication with the native heart of the patient using the system, to thereby receive signals corresponding to physiological parameters from the native heart.
 59. The system of claim 58, wherein the at least one hydraulic pumping assembly is two hydraulic pumping assemblies and the L-VAD and the R-VAD are each in fluid communication with a separate one of the two hydraulic pumping assemblies.
 60. A system for completely replacing cardiac ventricular function in a diseased native heart, the system comprising: a. a hydraulic pumping system; and b. a BI-VAD assembly having a L-VAD and a R-VAD, the L-VAD and the R-VAD having sufficient stroke volumes to supply the total cardiac blood flow output for the diseased native heart of a patient using the system, the L-VAD being disposed to at least partly replace the diseased left ventricle of the native heart of the patient and the R-VAD being disposed in normal use position to at least partly replace the diseased right ventricle of the native heart of the patient, with the inlet of the R-VAD being grafted to an artificial heart valve and the outlet of the R-VAD being grafted into the pulmonary trunk of the patient; wherein the L-VAD comprises: a. a first magnet having an open center and formed of high ferromagnetic-constant material; b. a first vessel surrounding the first magnet and defining a space in fluid communication with the aortic arch of the patient using the device, the first magnet being movable within the first vessel in substantially fluid-tight relation thereto; c. a second magnet formed of high ferromagnetic-constant material and being in magnetic communication with the first magnet, so that the first magnet and the second magnet are biased toward and tend to lock to one another, to thereby move in the same direction as one another; d. a second vessel encasing the second magnet and defining a space, the second magnet being movable within the space in substantially fluid-tight relation to the second vessel, the space defined by the second vessel being in fluid communication with a hydraulic pump for actuating the second magnet; and e. an one-way valve connected to the corresponding first magnet of the L-VAD , the one-way valve being movable with the first magnet of the L-VAD and closed when moving in a direction toward the aortic arch when the device is in normal use position in a patient using the device, to thereby cause blood of the patient to push through the aortic arch as the first magnet moves toward the aortic arch of the patient when the second magnet is actuated to move toward the aortic arch, the one-way valve in the L-VAD further being open when moving away from the aortic arch, to-thereby permit blood of the patient to flow through the one-way valve of the L-VAD into the space defined by the first vessel when the second magnet of the L-VAD is moved away from the aortic arch; and wherein the R-VAD comprises: a. a first magnet having an open center and formed of high ferromagnetic-constant material; b. a first vessel surrounding the first magnet and defining a space in fluid communication with the bifurcation of the pulmonary arteries of the patient using the device, the first magnet being movable within the first vessel in substantially fluid-tight relation thereto; c. a second magnet formed of high ferromagnetic-constant material and being in magnetic communication with the first magnet, so that the respective magnetic fluxes of the first magnet and the second magnet affect each other, so that the first magnet and the second magnet are biased toward and tend to lock to one another, to thereby move in the same direction as one another; d. a second vessel encasing the second magnet and defining a space, the second magnet being movable within the space in substantially fluid-tight relation to the second vessel, the space defined by the second vessel being in fluid communication with a hydraulic pump for actuating the second magnet; and e. an one-way valve being movable with the first magnet of the R-VAD and closed when moving toward the bifurcation of the pulmonary arteries when the R-VAD is in normal use position in a patient using the assembly, to thereby cause blood of the patient to push through the bifurcation of the pulmonary arteries as the first magnet of the R-VAD moves toward such bifuircation when the second magnet of the R-VAD is actuated to move toward the bifurcation, the one-way valve of the R-VAD further being open when moving away from the bifurcation of the pulmonary arteries, to thereby permit blood of the patient to flow through the one-way valve of the R-VAD into the space defined by the first vessel when the second magnet of the R-VAD is moved away from the bifurcation of the pulmonary arteries; wherein the at least one hydraulic pumping assembly comprises: a. a hydraulic pump having: an encapsulated pumping chamber for retaining hydraulic fluid therein, the pumping chamber having opposed first and second ends; at least one electromagnetic coil surrounding the pumping chamber; and a substantially solid high ferromagnetic-constant magnet disposed longitudinally, slideably and reciprocally within the pumping chamber to act as a piston for driving hydraulic fluid within the pumping chamber in response to signals from a battery/controller assembly; b. a fluid line having a first end and a second end, the first end of the fluid line being connected to and in fluid communication with an the first end of the pumping chamber and the second end of the fluid line being connected to and in fluid communication with the second end of the pumping chamber, the L-VAD and the R-VAD being in fluid communication with the fluid line at a point on the fluid line after the point of connection of the check valve and before the connection of the second end of the fluid line and the second end pump chamber; and c. a battery/controller assembly operatively connected to the check valve and to the at least one electromagnetic coil to provide electric power and control signals to the pump, the battery controller assembly in electrical communication with the native heart of the patient using the system, to thereby receive signals corresponding to physiological parameters from the native heart.
 61. A system for assisting cardiac ventricular function, the system comprising: a. a ventricular assist device (VAD) having: an open-centered magnet, at least one encapsulated electromagnetic coil in magnetic communication with the open-centered magnet; to thereby drive the magnet; and a one-way valve connected to the open-centered magnet, the one-way valve being movable with the magnet, and adapted to cause movement of blood from the diseased ventricle to and into the great vessel associated with the diseased ventricle; b. a battery/controller assembly operatively connected to the at least one electro-magnetic coil for energizing same and connected to the sino-atrial node of the patient when the system is in normal operative position in the patient to thereby provide signals to the VAD from the sino-atrial node to activate the at least one electromagnetic coil to optimally complement the function of the diseased ventricle of the patient's native heart. 